Injectable Biodegradable Polycaprolactone–SebacicAcid Gels for Bone Tissue Engineering

Christiane L. Salgado, Ph.D.,1,2 Elisabete M.S. Sanchez, Ph.D.,3 Cecılia A.C. Zavaglia, Ph.D.,3

Ana Beatriz Almeida, Ph.D.,4 and Pedro L. Granja, Ph.D.3,4

Tissue engineering constitutes a promising alternative technology to transplantation medicine by creating viablesubstitutes for failing tissues or organs. The ability to manipulate and reconstitute tissue function has tremen-dous clinical implications and will most likely play a key role in cell and gene therapies in the coming years. Inthe present work, a novel injectable and biodegradable biomaterial is reported that could be injected on thehuman body with a surgical syringe. The material prepared is a blend of polycaprolactone (PCL), a biode-gradable and elastic biomedical polymer, and sebacic acid, a natural polymer part of castor oil with lowmolecular weight to accelerate the slow degradation rate of PCL. The biocompatibility of the blend was eval-uated in vitro and its in vivo behavior was also assessed through subcutaneous and bone implantation in rats toevaluate its tissue-forming ability and degradation rate. The results allowed the conclusion that the gel isbiocompatible, promotes the differentiation of mesenchymal stem cells, and presents an adequate degradationrate for use in bone tissue engineering. In vivo the gel blends promoted tissue regeneration and adverse reactionswere not observed on subcutaneous and bone implants.

Introduction

Injectable biomaterials are becoming increasingly in-teresting and could constitute a great advancement in

many applications of controlled drug delivery and tissueengineering, largely due to their minimally invasive nature.Injectable systems in drug delivery can be used both forparenteral drug delivery or localized injection to an affectedsite. For the patient, injectable systems offer the advantage ofavoiding highly invasive surgical procedures and minimizepotential complications such as infections due to less expo-sure. Injectable scaffolds provide the ability to adapt to theshape of the defect cavity in which they are placed and canthus fulfill irregular defects. Also, they facilitate the incor-poration of cells and growth factors during the preparationof injectable solutions.1–5

For tissue engineering applications a scaffold materialmust be biocompatible. This applies not only to the materialitself, but also to possible degradation products in the case ofa biodegradable material.6 Another parameter that has to beconsidered in the design of a tissue engineering scaffold isporosity. Empty space within the scaffold network is con-sidered necessary to allow for tissue growth and diffusion ofnutrients to and waste products from the cells. Besides the

total porosity, pore size and interconnectivity are also im-portant variables.7 Ideally, the scaffold should be bioactiveand able to promote and guide cell proliferation, differenti-ation, and tissue growth. This could be achieved with theaddition of growth factors and the functionalization of thescaffold with proteins or adhesion-specific peptide sequencesthat often mimic the natural extracellular matrix and canprovide appropriate signals to cells.1,3,8,9 Two additionalfactors that have to be taken into consideration to match aspecific application are the mechanical properties and thebiodegradability of the scaffold. To be used for bone tissueengineering, for example, a scaffold must possess sufficientmechanical integrity to support the tissue, particularly dur-ing the early stages of growth.10 Moreover, the scaffoldshould degrade at a rate which ideally should match newtissue formation. A fast degrading scaffold that enables therelease bioactive agents and extracellular matrix componentsis desirable in some cases.11 Temporary devices are used onthose circumstances in which the natural tissue had beenweakened by disease, injury, or surgery and require a sup-port to reconstruct the tissue. As a consequence, the de-gradable implant would provide temporary mechanicalsupport, although it has to work properly by providing agradual transference of stress.11,12

1INEB-Instituto de Engenharia Biomedica, Universidade do Porto, Porto, Portugal.2Faculdade de Engenharia, Universidade do Porto (FEUP), Porto, Portugal.3School of Mechanical Engineering, State University of Campinas (UNICAMP)–R. Mendeleyev, Campinas, Brazil.4Sociedade Brasileira de Pesquisa e Assistencia para a Reabilitacao Craniofacial (SOBRAPAR), Campinas, Brazil.

TISSUE ENGINEERING: Part AVolume 18, Numbers 1 and 2, 2012ª Mary Ann Liebert, Inc.DOI: 10.1089/ten.tea.2011.0294

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The most widely used devices are degradable polymerssuch polyglycolic acid and polylactic acid (PLA) since theyprovide important advantages such as biocompatibility as-sociated with good mechanical properties, although theymay promote tissue necrosis due to local acidification andare expensive.13 Another biodegradable polymer used asbiomaterial is the polycaprolactone (PCL). PCL had beenproposed for varied biomedical applications, includingscaffolds for tissue engineering of bone and cartilage, owingto its unique elastic properties among the biodegradablepolymers.14 Due to its relatively low melting point, PCL maybe easily processed by conventional procedures.15 To im-prove its mechanical properties, PCL has been blended orcopolymerized with other polymers, such as PLA or PLGA.16

However, and despite its many interesting properties thebiodegradation of PCL occurs at a slow rate in vivo, it couldbe > 6 months and lower than 2 years, limiting its utilizationas tissue engineering scaffold.17–19

In an attempt to improve the in vivo biodegradation rate ofPCL, the present work reports the development of an in-jectable PCL-based gel blended with a biodegradable naturalpolymer, sebacic acid (SA), which degrades faster than purePCL without loss of mechanical resistance.

Materials and Methods

Materials preparation

The PCL gel-like matrices were prepared using 1.5 g ofPCL (Mw*80.000, from Sigma-Aldrich) dissolved in 5 mL ofdichloromethane (Merck) and 2.5 mL of polycaprolactonediol (Solvay caprolactone) with or without SA (5%, 100 mg)(Merck) stirred on 20 mL of calcium chloride solution(0.01 g/mL, from Synth), at 60�C. The material was thenwashed for 30 min with 10 mL dimethylsulfoxide (Merck),for 3 h with 50 mL phosphate-buffered saline (PBS, fromSigma-Aldrich), and sterilized under UV light (UV-C,253.7 nm, 30 W; Philips) for 20 min.20 The crosslinking of thepolymer gel starts with the synthesis temperature (60�C) andends after 24 h in vitro (simulated body fluid solution at37�C), by solvent exchange.

Biological characterization

Cell culture. In this study human mesenchymal stemcells (hMSCs, from American Type Culture Collection) wereused. hMSCs were expanded in MSC growth medium(Sigma-Aldrich) with 10% fetal bovine serum (Cultilab) untilthe sixth passage, in a humidified atmosphere at 37�C and5% CO2. After reaching confluence (90%), cells were culturedwith materials at a density of 2 · 103 cells/mL. Osteoblasticdifferentiation of hMSCs was stimulated by culturingthe cells in osteogenic media for 24 h, using dexamethasone(10 - 8 M; Sigma-Aldrich), ascorbic acid (30mg/mL; Sigma-Aldrich), and beta-glycerophosphate (10 mM; Sigma-Aldrich).In all culture media 0.5% of a solution containing streptomycin(3 · 10 - 4 M; Sigma-Aldrich) and penicillin (5 · 10 - 4 M; Sigma-Aldrich) was added.

In vitro cell viability. The hMSC viability was accessed bythe alamar blue (Invitrogen) assay.21 As a preliminary eval-uation, the cells were seeded on the materials surface and,after 2 and 7 days, the culture medium was removed, and the

samples were washed with PBS and then incubated in 500mLof alamar blue solution for 4 h. Afterward, 100mL of thereactive solution was transferred to a 96-well black plate(Dow Corning). The fluorescence were measured in a platereader (Spectra Max Gemini XS) at lex = 530 nm and lem =590 nm.

In vitro differentiation studies. Alkaline phosphatase(ALP) and total protein quantification: To analyze early celldifferentiation, hMSCs were cultured on the material surfacefor 1, 3, and 10 days and the ALP activity was measured.After cell culture as previously described the cells were re-moved from the materials surface with trypsin solution(0.5%) and then lyzed with Triton X-100 solution (10%;Sigma-Aldrich). The p-nitrophenol assay was used, in whichthe colorless nitrophenyl phosphate disodium salt (Sigma) ishydrolyzed by the ALP (Sigma) at 37�C and pH 10.5 to formyellow p-nitrophenol. After 1 h the chemical reaction wasstopped with NaOH (0.02 M; Synth) and the optical absor-bance was measured on a plate reader (405 nm). The stan-dard curve was obtained with 10mg/mL p-nitrophenol(Sigma).22 Total protein content was determined using theBio-Rad Protein Assay (DC� Protein Assay; BioRad), ac-cording to the manufacturer’s indications (BCA reagent,calcium bicarbonate, sodium tartarate, copper sulfate, andbicinconic acid) for 1 h at 37�C and the absorbance read at655 nm. The calibration curve was done with an increasingconcentration of serum bovine albumin solution (1–0.1 g/mLin PBS; Sigma).

von Kossa assay: The presence of phosphate on hMSCsseeded on the biomaterial surface (2 · 105 cells/mL) wasevaluated using the von Kossa assay after 10 and 15 days inosteogenic medium changed every other day. Afterward, thesamples were washed twice with PBS and fixed with 4%formaldehyde (Merck) for 10 min. Then, the materials werewashed three times with distilled water and incubated with2% AgNO3 solution (Sigma-Aldrich) for 30 min, at roomtemperature, under UV light (UV-C, 253.7 nm, 30 W; Philips).The materials were then washed three times with PBS andsodium thiosulfate solution (Sigma-Aldrich) was added, for5 min. Samples were washed again with PBS and let dry atroom temperature. The sodium thiosulfate reacted with thesilver nitrate, and precipitated on the calcium in the extra-cellular matrix. The materials were observed in an opticalmicroscope (Olympus model GX51).

Microscopical analyses. Scanning electron microscopy:After culturing the cells on the materials surface for 10 days,the culture medium was removed and the materials washedthree times with PBS. The samples were prepared for scan-ning electron microscopy (SEM) ( Jeol JXA 840 A, 40 me and10 kV) observations by adding 1.5% of glutaraldehyde and0.14 M sodium cacodylate for 30 min. Later, the sampleswere washed with distilled water and then dehydrated witha sequence of ethanol solutions (70%–100%) for 10 min eachand finally dried at room temperature for 24 h.23 Sampleswere gold-sputtered using the Balzers equipment (modelSCD 050, USA) and observed at 20 kV.

Confocal laser scanning microscopy: For confocal laserscanning microscopy (CLSM) observations the cells werecounted, seeded on the biomaterials surface, and cultured inwells for 3 and 10 days. After that time the culture medium

138 SALGADO ET AL.

was removed, the samples were washed three times withPBS, and formaldehyde at 3.7% was added for 10 min. Thecells were then treated with Triton X-100 solution (0.1%) for5 min and then washed three times with PBS solution. Fi-nally, 400 mL of Alexafluor 488 (Invitrogen) was added for60 min, followed by washing three times with PBS, and400 mL of propidium iodide (Sigma-Aldrich) were added for5 min. The materials were washed with PBS before obser-vation under a Confocal Laser Scanning Microscope (LeicaTCS STED CW).

In vivo animal implantation studies. Male Wistar rats(150–250 g; Bioterio UNICAMP) were used for in vivo eval-uation according to the international rules submitted andapproved by the Animal Ethical Committee of the BiologyFaculty of Campinas State University, process number CEEA1534-2. The materials were implanted with a subcutaneousand bone (tibia) injection of the PCL-SA blends. The materialwas injected with a common syringe with a 0.05 mm needle.The negative control was a PBS solution injected the sameway as the polymer gel. Five animals were used for eachgroup of periods of 30, 60, and 90 days. Before surgery theanimals were sedated with an intramuscular injection ofketamine (0.5 mL/300 g weight; Dopalen�) and xilasinechloride (0.2 mL/300 g weight; Kensol�). The animals wereshaved and the area of implantation was disinfected withethanol (Synth) at 70%.

The subcutaneous implants were injected (0.5 mL) with asyringe (Injekt�) and a hypodermic needle (0.05 mm; Steri-can�) under the skin on the animal left dorsal lumbar region.The PBS solution was injected with the same volume on theright side of the animal. The bone defect was prepared on thetibia with a drill of low rotation (spheroid with 4 mm, Car-bidePM014; Wilcos), and was irrigated with NaCl (Synth)sterile solution (0.09%) to remove blood clots. The defect sizewas limited by the dimension of the tibia, so it was not acritical bone defect that should be between 5 and 8 mm forthese animals. After the defect was clear, the gel was injected,the same way as the subcutaneous implant, inside the bonedefect. The wound was sutured with a 4-0 nylon suture(Carpule, Nipro). The animals were maintained in boxeswith food and water at libitum. They were sacrificed by neckdislocation and the tibia were removed and the implant areacut with a carbon disc and immersed into formaldehyde(Merck) solution (10%) overnight. The formaldehyde solu-tion was changed by an ethanol solution (70%) for the sub-

cutaneous implants and decalcification solution (EDTA-0.7 g;potassium and sodium tartarate-8 g; sodium tartarate-0.14 g;HCl-120 mL; distilled water-880 mL, all from Synth) for thebone tissue. The bone defect area with the material wasprocessed for histological resin embedding after dehydrationin ascending concentration of ethanol and xylene. Aftercompleting the embedding process the paraffin blocks weretrimmed and sawed to enable sections of 5mm in thickness.The samples were mounted onto microscope slides andstained using hematoxylin and eosin solution (Synth). Theanalyses were carried out by optical microscopy (Olympusmodel GX51).

Statistical analyses. Data are presented as mean –standard deviation (n = 3) and were analyzed using the t test.Differences between groups were considered statisticallydifferent when p < 0.05.

Results and Discussion

A novel injectable, biodegradable PCL-based material in-corporating SA was obtained. This material was prepared inthe form of an injectable and in situ crosslinking gel, thusallowing its use in minimally invasive surgical procedures.The polymer gel prepared is a transparent viscous liquidafter exposure to water, exhibiting a gel-like behavior. Thepolymer material separated completely from the calciumsolution, and when PCL-SA gel was exposed to air the sol-vent evaporated and the viscous liquid became a rigid gelafter 1 h (Fig. 1A–C). In vivo, the solvent in the polymer isprobably exchanged by the blood clotting fluid and thematerial slowly hardens in about 24 h.

The toxicity of organic products usually used to preparesynthetic polymeric materials constitutes a limiting factor fortheir use as implant materials. The PCL-SA blend was shownto be cytocompatible on the direct contact assay usinghMSCs. Figure 2 shows the results of cell metabolic activityby the alamar blue assay. These results are presented asrelative percentage compared with the negative control (tis-sue culture polystyrene, TCPS, from culture wells), withoutany material (100%). Cell metabolic activity was calculatedusing the following equation:

Cell viability (%)¼ (Nt=Nc) · 100, ð1Þ

where Nt is the optical density of the sample group and Nc isoptical density of the negative control group (TCPS).

FIG. 1. (A) Polycaprolactone sebacic acid (PCL-SA) viscous gel sample, as prepared; (B) image showing the gel-likebehavior of PCL-SA; (C) after 1 h exposure to air, the PCL-SA gel becomes a solid stable gel. Color images available online atwww.liebertonline.com/tea

POLYCAPROLACTONE–SEBACIC ACID GELS IN TISSUE ENGINEERING 139

Figure 2 presents hMSC viability in contact with test sam-ples, showing cell viability over 90% for all the materials tested(PCL alone and PCL-SA), comparing to the negative control(TCPS), using the alamar blue assay. The PCL-SA presented asimilar decrease in cell viability as PCL after 7 days, althoughpresenting cell viability always higher than TCPS.

hMSC cells were cultured on the surface of PCL-SA geldiscs during 3 and 10 days. After the incubation period thesamples were analyzed by SEM and CLSM (Figs. 3 and 4).SEM observations after 10 days (Fig. 3) indicated that hMSCadhered, spread, and interconnected and also showed nor-mal morphology on the gel surface. However, these cellsexhibited a lower proliferation rate than the control, whichdid not allow monolayer coverage of the whole surface.

The samples observed under confocal microscopy showeda few adhered cells on the initial culture times (Fig. 4A). At10 days colonies of spread cells were observed all over thegel surface (Fig. 4B). This morphology is a characteristic ofhMSC cells well adhered to a surface.

The hMSCs adhered to the PCL-SA gel after 24 h, whenthe culture medium was changed to the osteogenic medium

to promote their differentiation. At determined time pointsthe cells were lyzed with Triton X-100 solution (10%) and theALP activity was measured with the p-nitrophenyl assay.The total protein amount was also measured and Figure 5shows the ALP concentration as a function of the total pro-tein concentration. At all the times of incubation (1, 7, and 14days) the ALP activity in the presence of the PCL-SA gel waslower than that of the control (PCL), although this differencewas minimal at the latest time point.

A late marker of bone cell differentiation is the depositionof calcium phosphate on the extracellular matrix. This phe-nomenon was analyzed using the von Kossa assay. Figure 6shows phosphate deposits resulting from the cells culturedon PCL-SA gel surface, which appear as characteristic darknodules (silver nitrate). The cells were also seeded on purePCL and the formation of calcium nodules was not observed.This fact may indicate that cells do not mineralize on PCL ormay indicate an earlier differentiation in PCL-SA. The min-eralization in PCL-SA was most likely favored by the pres-ence of calcium in the gel preparation, which can act as anagent of nucleation of mineralization. In a study using hu-man marrow stromal cells seeded into PCL scaffolds, anaccomplished proliferation was observed, although the dif-ferentiation was hardly observed.24 Therefore, the PCL-SAgel presented here, with improved biodegradability and ca-pable of supporting or promoting proliferation and differ-entiation of mesenchymal cells, shows an improved ability toinfluence cell behavior, which is of paramount importance intissue engineering applications.25

The in vivo biocompatibility evaluation of the PCL-SA gelrevealed further features than the in vitro analyses. The ma-terial was injected subcutaneously and also in a bone site.The solvents used during the preparation of the histologicalsections resulted in the PCL-SA material being dissolvedaway during the polymer embedding; thus, gel scaffolds arerepresented in Figures 7–9 as empty lacunae surrounded byconnective tissue, manually; the reduction of the implantarea was observed on different time points.

After 30 days of implantation (Fig. 7A, B) the tissueformed inside the material had normal connective tissuemorphology with the presence of fibroblasts, blood vessels,and mononuclear cells. In the surrounding tissue small cellswith a central dense nucleus, corresponding to the morphology

FIG. 2. Alamar blue assay of human mesenchymal stemcells (hMSCs) cultured on PCL and PCL-SA blend for 2 and 7days.

FIG. 3. (A) Scanning electron microscopy micrographs of hMSCs cultured on PCL-SA for 10 days. (B) Amplification of (A).

140 SALGADO ET AL.

of monocyte inflammatory cells, could be observed. This pro-cess could be induced by the material cross-linking. A reducedchronic inflammation could be observed, as shown by thepresence of some inflammatory cells, such as macrophages andlymphocytes at the tissue formed inside the implant, and thismild cell response was most likely fundamental for the tissueregeneration. Independent studies with PCL-based scaffoldsrevealed, after 3 months of subcutaneous and intramuscularimplantation, a high concentration of adipose-like cells insidethe pores. After 6 months the implants showed fibrous tissuereplacing the adipose-like tissue, along with higher number ofblood vessels.19

The PCL-SA gel surface hydrophilicity was probably im-proved due to the presence of SA on the top layer, whichadditionally provided partial solubility in water. Also, giventhe presence of calcium used in the cross-linking, a numberof positively charged particles were available near the sur-

face. The inflammatory response to a biomaterial is mediatedby a complex interaction between macrophages, giant cells,and cytokines.26 The chemical properties of the surface areknown to influence the release of cytokines (e.g., interleukin[IL]-1b, IL-6, IL-8, and tumor necrosis factor-a) and the ac-tivation and adherence of macrophages during the foreignbody response.27,28 In hydrophilic materials with a positivecharge, the attraction and migration of macrophages hasbeen reported as lower, and the formation of giant cells in-hibited, comparing to hydrophobic or neutral materials.19

Hence, the present materials would probably obstruct ad-hesion of macrophages and giant cells formation, which, inturn, would promote fibroblasts migration and collagenformation (fibrous capsule). A thin layer of collagen fiberswas also observed surrounding the material. The new tissueformed was vascularized with blood vessels evidencingblood cells (Fig. 7B). Some degradation of the PCL-SA gelcould already be observed even after 30 days of implantation(Fig. 7A).

After 60 and 90 days of implantation (Fig. 8A, B, respec-tively) just a few fragments of the material could be found,with connective tissue growing inside the implant and re-placing the original space filled by the gel. Some of the PCL-SA gel particles were surrounded by collagen fibers (fibrouscapsule). Connective tissue was formed, with collagen fibersand few inflammatory cells (macrophages and lymphocytes).The new tissue formed showed the presence of new bloodvessels (angiogenesis) with blood cells. Since the vasculari-zation of an implant is perceived as a limiting factor fortissue regeneration inside an implant material,29 the presencePCL-SA gel provides clear indication of an adequate in vivobehavior, with formation of new blood vessels. At 90 days ofimplantation (Fig. 8B) the implanted gel was reduced tosmall polymer fragments, showing an advanced stage ofdegradation. Probably, these small portions remained un-degraded because they are most likely constituted of highlycrystalline PCL fragments. A collagen fibrous capsule sur-rounded these remaining fragments, although the connec-tive tissue formed inside the material presented normal

FIG. 4. Confocal laser scanning microscopy images of hMSC cultured over PCL/SA gel surface for (A) 3 days and (B) 10days. Color images available online at www.liebertonline.com/tea

FIG. 5. ALP activity of hMSCs cultured on PCL and PCL-SAblend surface for 1, 7, and 15 days. ALP, alkaline phosphatase.

POLYCAPROLACTONE–SEBACIC ACID GELS IN TISSUE ENGINEERING 141

morphology. The implantation of the present PCL-SA gelrevealed no adverse response, such as tissue necrosis, abscessformation, calcification, or tumor induction. The present re-sults suggest adequate overall biocompatibility of the mate-rial after subcutaneous implantation.

The PCL degradation depends highly on the polymermolecular weight. Higher molecular weight means longerchain length of the material and, consequently, the degra-dation takes longer.30 Initially, the in vivo degradation of PCLoccurs as a nonenzymatic bulk hydrolysis, resulting in anearly transient inflammatory response for the first 2 weeks ofimplantation.31 PCL’s molecular weight may decrease after 9months and small fragments of the material should be de-tected on the implant area, typical of a resorbable polymer.32

In a study of PCL in vivo degradation, low molecular weightpolymer powder was used. It was observed that the materialrapidly degraded and after 13 days hydroxyl caproic acid(metabolite of PCL after degradation) could be observedinside macrophages and giant cells.32

Jabbari et al. reported similar small inflammatory pro-cesses after in vivo studies using an injectable matrix of PCLand fumarate.33 In another study with PCL and differentpolymers that were evaluated in vivo for biocompatibility

analyses, it was shown similar results of implanted materi-als.34 In another study a copolymer of PCL and SA and ri-cinoleic acid was evaluated in vivo for its controlled drugrelease ability.35 The histological analyses showed the pres-ence of acute inflammation after 1 week of implantation.However, inflammatory cells were not observed after 1month.35 Another study of the same group investigated theimplantation of an injectable copolymer of PCL and ricinoleicacid, which cross-linked in vivo by solvent exchange. Theimplant stayed in the same place after 24 h, just like when oilis injected in the subcutaneous space.36 In a study of Modiet al., an implanted PCL-PLA copolymer showed an acuteinflammatory response right after the injection of the bio-material.37

The bone implantation of the PCL-SA gel for 30, 60, and 90days revealed some degree of continuous degradation al-though slower than after subcutaneous implantation. Thebone defect was totally filled with the injected polymer, butas the bone marrow has blood cells, a blood clot formedunder the polymer hindering full contact with the bone tis-sue. The constant bleeding resulting from the procedure alsoconstitutes an obstacle to an adequate contact on the inter-face between the gel and the surrounding tissue. As part of

FIG. 6. Optical microscopy images of hMSC cultured on PCL-SA gel, as evaluated by the von Kossa assay (circles) after (A)10 days and (B) 15 days of culture. Amplification of 40 · .

FIG. 7. Optical microscopy images of histological sections of PCL-SA gel after 30 days of subcutaneous implantation. Thetissue was processed and stained with hematoxylin and eosin (H&E). The PCL-SA gel particles can be seen as empty spacesinside the tissue. Image amplifications of (A) 40 · and (B) 100 · . BV, blood vessels. Color images available online atwww.liebertonline.com/tea

142 SALGADO ET AL.

the normal healing response, probably the gel promotedplatelet aggregation and inflammatory cells migration. Thisblood clot could promote cell migration due to growth fac-tors secretion into the wound environment.38 After 30 days,the new tissue formed was mineralized, although disorga-nized, with globular aspect (Fig. 9A). This structure did notshow presence of Haversian canals (series of tubes around

narrow channels formed by lamellae)39 and the bone for-mation is completely disorganized. The new bone formedhad similar aspect to the negative control, which was thedefect without any material, the bone could regenerate alone(Fig. 9C). As explained before the size of the animals tibia didnot allow creating a critical defect without hindering theanimal’s movement. A thin fibrous capsule was observed

FIG. 8. Optical microscopy images of histological sections of PCL-SA gel after subcutaneous implantation. The tissue wasprocessed and stained with H&E. The PCL-SA gel particles could be seen as empty spaces inside the tissue. Materialimplanted after (A) 60 and (B) 90 days. Image amplifications of (A) 100 · and (B) 40 · . Color images available online atwww.liebertonline.com/tea

FIG. 9. PCL-SA gel implanted in rat tibia after (A) 30 days, (B) 60 days, (C) 90 days, and (D) negative control (phosphate-buffered saline) after 90 days of surgery. The PCL-SA presence is shown with arrows. Image amplifications of 40 · . BD, bonedefect; BM, bone marrow. Color images available online at www.liebertonline.com/tea

POLYCAPROLACTONE–SEBACIC ACID GELS IN TISSUE ENGINEERING 143

involving the PCL-SA gel (shown with arrows in Fig. 9A)and small blood vessels were formed. The presence of oste-oblasts and osteoclasts was observed at higher amplification(Fig. 10A), probably for remodeling the new bone tissue (Fig.9A). After 60 days of implantation, the bone formed showeda more organized structure, similar to lamellae. This struc-ture is characterized by the presence of Haversian canals,aligned in a longitudinal direction with compact and densemorphology (Fig. 9B). Some gel particles still remain, sur-rounded by disorganized connective tissue (shown with ar-rows in Fig. 9B). As previously observed, osteoblasts andosteoclasts were found on the surface of new bone tissue. Thecells could only be observed and identified at higher am-plification (30 and 60 days of implantation—Fig. 10A, B). Inthe negative control (PBS solution) bone tissue formationwas delayed compared to the polymer gel. The structure wasmore globular and disorganized (Fig. 9D). The implants weredifficult to identify in the histological sections due to materialdegradation as the tissue was formed. During the histologi-cal processing the remaining materials may also been re-moved.

After 90 days of implantation the bone tissue formed wascompact and organized, and the presence of osteoclasts atthe surface was not observed (Fig. 10C). Still, a thin fibrouscapsule could be observed surrounding the implanted bio-material (Fig. 9C). In studies carried out by Sawyer and co-workers40 implant materials based on PCL and tricalciumphosphate, with or without collagen, had been implanted in

rat calvaria with bone morphogenetic protein 2. After 4weeks of implantation the results reported were similar tothe ones presented here using PCL-SA gel, where a disor-ganized bone tissue formation was found. After 15 weeks thedefect was completely closed, with the formation of a moreorganized bone tissue. However, since the defect was not ofcritical size, the closure of the defect was not observed at alltime points, when no materials were implanted.40

Concerning the in vivo biodegradation of materials theimplanted gel remained macroscopically unchanged in termsof morphology after the different times of implantation at abone site (30, 60, and 90 days). Nevertheless, an improvedtissue formation and a more organized bone structure wereobserved along the time of implantation (Fig. 9A–C). Simi-larly, in an independent study, after 2 years, implants of aPCL scaffold in rabbit calvaria showed bone formation sur-rounding the material, but the polymer structure still re-mained in the tissue and the bone defect was not completelyregenerated.19

Conclusions

PCL-SA gel, capable of crosslinking in vivo, is here re-ported for the first time with the purpose of developing in-jectable biomaterials for minimally invasive therapeuticprocedures. In contact with MSCs this gel was shown to becytocompatible, and able to support cells metabolic activityand maintain their viability. Cells cultured on crosslinked

FIG. 10. PCL-SA gel implanted in rat tibia after (A) 30 days, (B) 60 days, and (C) 90 days of surgery. The PCL-SA presence isshown with arrows. Image amplifications of 100 · . OC, osteoclast; OB, osteoblast. Color images available online atwww.liebertonline.com/tea

144 SALGADO ET AL.

samples showed viability higher than 90% and were able toattach, proliferate, and differentiate over the material surface,as shown by determination of the ALP activity and assess-ment of mineralization. In vivo subcutaneous implantationstudies showed a reduced inflammatory response with newvascular tissue formation inside and around the implant anda considerable biodegradation of the material after 90 days ofimplantation. Implantation in rat tibia showed formation ofnew bone tissue under the implants in all the time pointsanalyzed.

Acknowledgments

Authors are grateful to Coordenacao de Aperfeicoamentode Pessoal de Nıvel Superior (CAPES-4633-06-7) and Con-selho Nacional de Desenvolvimento Cientıfico e Tecnologico(CNPq-142956/2005-8) for funding and the scholarship forDr. Salgado.

Disclosure Statement

No competing financial interests exist.

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Address correspondence to:Christiane L. Salgado, Ph.D.

Faculdade de EngenhariaUniversidade do Porto

DEMMRua Dr. Roberto Frias, s/n

Porto 4200-465Portugal

E-mail: chris@fe.up.pt

Received: May 20, 2011Accepted: July 25, 2011

Online Publication Date: September 14, 2011

146 SALGADO ET AL.