Three-Dimensional Poly (ε-caprolactone) Bioactive Scaffolds with

Oct 2, 2012 - aminolysis. Causa et al.30 also proposed a systematic study of peptide ligand organization and spatial distribution on PCL surfaces, eva...
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Three-Dimensional Poly(ε-caprolactone) Bioactive Scaffolds with Controlled Structural and Surface Properties A. Gloria,*,† F. Causa,‡ T. Russo,† E. Battista,‡ R. Della Moglie,‡ S. Zeppetelli,† R. De Santis,† P. A. Netti,‡,§ and L. Ambrosio†,§ †

Institute of Composite and Biomedical Materials, National Research Council, P.le Tecchio 80, 80125, Naples, Italy Interdisciplinary Research Centre on Biomaterials (CRIB), University of Naples “Federico II”, and Center for Advanced Biomaterials for Healthcare (CABHC), Istituto Italiano di Tecnologia (IIT), P.le Tecchio 80, 80125, Naples, Italy

‡

S Supporting Information *

ABSTRACT: The requirement of a multifunctional scaffold for tissue engineering capable to offer at the same time tunable structural properties and bioactive interface is still unpaired. Here we present three-dimensional (3D) biodegradable polymeric (PCL) scaffolds with controlled morphology, macro-, micro-, and nano-mechanical performances endowed with bioactive moieties (RGD peptides) at the surface. Such result was obtained by a combination of rapid prototyping (e.g., 3D fiber deposition) and surface treatment approach (aminolysis followed by peptide coupling). By properly designing process conditions, a control over the mechanical and biological performances of the structure was achieved with a capability to tune the value of compressive modulus (in the range of 60−90 MPa, depending on the specific lay-down pattern). The macromechanical behavior of the proposed scaffolds was not affected by surface treatment preserving bulk properties, while a reduction of hardness from 0.50−0.27 GPa to 0.1−0.03 GPa was obtained. The penetration depth of the chemical treatment was determined by nanoindentation measurements and confocal microscopy. The efficacy of both functionalization and the following bioactivation was monitored by analytically quantifying functional groups and/or peptides at the interface. NIH3T3 fibroblast adhesion studies evidenced that cell attachment was improved, suggesting a correct presentation of the peptide. Accordingly, the present work mainly focuses on the effect of the surface modification on the mechanical and functional performances of the scaffolds, also showing a morphological and analytical approach to study the functionalization/bioactivation treatment, the distribution of immobilized ligands, and the biological features.

1. INTRODUCTION The loss or failure of an organ or tissue represents a frequent and devastating problem in health care. Thus, the need for substitutes to replace or repair tissues or organs problems is overwhelming. For this reason, tissue engineering is becoming an important field of research.1−3 Basically, three approaches have been studied singularly or in combination in order to achieve tissue regeneration: cell-based therapies, tissue-inducing factors, and biocompatible scaffolds.3−9 In the field of tissue engineering, the design of a scaffold able to guide the process of tissue regeneration is one of the most challenging goals. The ideal scaffold should promote and control specific events at the cellular and tissue levels, as a results of its unique chemical, biochemical, and biophysical cues.7 Consequently, a scaffold has to satisfy several requirements:3,10,11 it should be made of highly biocompatible material that does not have the potential to elicit an immunological or clinically detectable foreign body reaction.10 Over the last two decades, the concept of cell guidance has also been progressively revised7−13 since a new knowledge about the complex features of cell-material interaction has come to light. Thus, novel scaffold materials based on the cell © 2012 American Chemical Society

guidance concept have been developed, benefiting from contemporary advances in the fields of molecular biology and materials science.3,14 Several polymeric and composite materials have been used to make three-dimensional (3D) porous scaffolds, using both conventional methods and more advanced manufacturing processes.10,11,15−24 Rapid prototyping is a common name for a group of techniques that can generate a physical model directly from computer-aided design (CAD) data and may act as a methodical interface between tissue and engineering.25−27 Differently from conventional processing techniques, rapid prototyping offers the possibility to strictly control pore geometry, size, and interconnectivity, as well as the spatial distribution of pores within the structure. Among novel rapid prototyping techniques for scaffold fabrication, 3D fiber deposition2,19−24 has recently emerged as a method to manufacture well-defined and custom-made scaffolds for tissue regeneration, with 100% interconnected pores. Received: May 25, 2012 Revised: October 2, 2012 Published: October 2, 2012 3510

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Three-dimensional (3D) fiber deposition represents a modified technique of 3D plotting15,16 for the extrusion of highly viscous polymers, and is essentially a fused deposition technique in which a molten polymer is extruded and deposited from a servo-mechanically controlled syringe applying pressure. With regard to materials, many biodegradable synthetic polymers such as polyglycolic acid (PGA), poly(lactic acid) (PLA), poly(lactide-co-glycotide) (PLGA) and poly(caprolactone) (PCL) have been already used for making scaffolds to support the regeneration of several tissueengineered organs.3,7 However, the poor cytocompatibility of the synthetic polymers leads to the inefficiency of the scaffold in obtaining a friendly interface with living cells. Consequently, in order to improve their cytocompatibility, modifications of the tissue engineering polymer-based materials are needed. In tissue engineering, cell adhesion to the scaffold surface is a critical factor since adhesion occurs before other biological events such as cell spreading, migration and differentiation. In particular, cell adhesion is strongly related to the surface properties of biomaterials, and it is influenced by substratum surface properties, such as surface charge, wettability, roughness, and topography. Most conventional materials do not meet the criteria for serving as tissue engineering scaffolds and, for this reason, many surface modification techniques have been developed to alter the surface properties of these materials.28−30 As the interaction between living cells and materials occurs mainly on the interfacial layer, many surface modification techniques such as γ-ray irradiation, plasma treatment, endgrafting, ozone oxidization, or in situ polymerization have been already considered to modify the materials surface properties, the aim being to improve the cytocompatibility of the polymeric materials without altering their bulk properties.28−30 However, PCL is a biodegradable aliphatic polyester that has been suggested for a wide range of applications such as drug delivery systems, tissue-engineered skin, scaffolds for supporting fibroblast and osteoblast growth.30−34 A previous study also highlighted that a dynamic seeding of osteoblasts and endothelial cells onto a 3D fiber-deposited PCL scaffold may represent a useful approach in order to achieve a functional hybrid in which angiogenesis, furnished by neo-vascular organization of endothelial cells, may further support osteoblasts growth.24 PCL molecules are characterized by a great number of ester groups (−COO−) that can be hydrolyzed to carboxylic acid under alkaline conditions. Moreover, amino groups can be introduced onto the polyester surface through a reaction with diamine, providing that one amino group reacts with the −COO− group to form a covalent bond, −CONH−, while the other amino group is unreacted and free.29 However, the hydroxylterminated chains will also be yielded on the polyester surface during this process. The decreasing of surface hydrophobicity, neutralization of the acid originated from the scaffold degradation, and the possibility to provide active sites through which other biomolecules such as gelatin, collagen, or arginine-glycine-aspartic acid (Arg-Gly-Asp or RGD) peptides can be immobilized, are the main advantages in tissue engineering that can be derived from the introduction of the amino groups.29 To understand the mechanisms of cell-material interaction, the structure of biological tissue has to be taken into consideration. Biological tissues consist of cells immersed in the extracellular matrix (ECM), which is a complex mixture of proteins and glycosaminoglycans with both mechanical and signaling functions. Integrin heterodimers bind to specific

amino acid sequences, such as the RGD recognition motif that is largely present in many ECM proteins.35 Thus, small synthetic peptides (a few hundred daltons) that contain the RGD amino acid sequence can mediate cell attachment as well as the large parental molecule (a hundred thousand daltons).30 Benefiting from this basis, great efforts have been made to develop biomimetic approaches for immobilizing short peptides (i.e., RGD) onto synthetic or natural substrates, the aim being to obtain multifunctional biomaterials able to promote and improve cell attachment.36,37 In this context, interesting quantitative results have also been obtained about cell spreading and formation of focal contacts. In particular, a minimum RGD density of 1.0 × 10−15 mol/cm2, that corresponds to a spacing of 140 nm between peptide ligands, seems to be sufficient to promote cell spreading, while a density of 1.0 × 10−14 mol/cm2 has been found to favor the formation of focal contacts.30,38 As PCL is a synthetic polymer, it does not possess molecular motifs for cell recognition.30,39 To promote cell adhesion, its backbone has to be suitably modified by introducing functional groups for the following RGD conjugation.40,41 Most of modification methods may show many problems such as low level of functional groups, lack of control of peptide immobilized on PCL surface, and the eventual presence of uncontrolled degradation products. In addition, the bioactive groups may not be covalently attached but only adsorbed onto the surface, thus leading to the possibility of being removed or exchanged upon introduction into in vitro culture or in vivo implantation. On the other hand, it has been demonstrated that chemical methods can be successfully used for bioactivating polymer surfaces. Functional groups have been introduced on PCL via hydrolysis,39,42 aminolysis,39 plasma treatment,43,44 or copolymerization.45 In the case of PCL, amine groups were introduced onto film surfaces through treatment with a diamine, before attaching peptide sequences (i.e., RGD) using either glutaraldehyde, carbodiimide, or epoxy-amine chemistry.30,46 Biological studies then evidenced an enhancement in cell adhesion and spreading on these modified surfaces,47 especially on PCL containing laminin-derived peptides sequences IKVAV (Ile-Lys-Val-AlaVa), YIGSR (Tyr-Ile-Gly-Ser-Arg), or integrin-derived peptides sequence RGD, covalently linked to the polymer surface. Santiago et al.47 used a two step-procedure to immobilize peptides onto the polymer surface, involving a treatment with 1,6-hexanediamine followed by the use of 1-ethyl-3-(dimethylaminopropyl) carbodiimide; while Taniguchi et al.48 also modified PCL with poly(ethylene oxide) grafts before coupling with RGD containing peptides, thus obtaining an improvement in cellular responses. Further recent works49,50 have highlighted the possibility to bioactivate 3D PCL scaffolds with RGD after aminolysis. Causa et al.30 also proposed a systematic study of peptide ligand organization and spatial distribution on PCL surfaces, evaluating the effective peptide distribution able to activate specific cell functions (i.e., adhesion or differentiation). In particular, they performed the grafting of the synthetic peptide Gly-Arg-GlyAsp-Tyr (GRGDY), which contains the RGD sequence of several adhesion molecules, onto PCL sheets using a two-step procedure similar to those already reported in the literature46,47 involving a polymer aminolysis to graft primary amines on the film surface and a subsequent conjugation of the RGD motif. Unlike other works, each step 3511

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Figure 1. Scheme of the two-step procedure used to graft GRGDY peptides to PCL fibers of the 3D scaffolds. DEA and IPA indicate 1,6hexanediamine and isopropanol, respectively. The chemical structure of DGDGE is also reported. Micro-Computed Tomography. Micro-computed tomography (Micro-CT) was performed through a SkyScan 1072 system (Aartselaar, Belgium) using a rotational step of 0.9° over an angle of 180°, in order to analyze the internal structure of the 3D scaffolds fabricated, pore shape, and size. Cross sections and 3D model of PCL scaffolds were reconstructed using Skyscan’s software package and Image J software, which allowed us to visualize and analyze the results from the Micro-CT system scan. The pore network was visualized and the pore interconnectivity was studied. Aminolysis of 3D PCL Scaffolds. A specific procedure, which was already shown in a previous work,30 was used to insert covalently amino groups onto the fiber surface of the 3D PCL scaffolds, using 1,6-hexanediamine (DEA). Briefly, 3D fiber-deposited scaffolds were immersed at different times in 0.08 g/mL DEA/isopropanol (IPA) solution at 37 °C. The aminolysis reaction was carried out in a custommade reactor thermostat in a water bath with adequate magnetic stirring for a suitable time, in batch mixer processing conditions. After the aminolysis treatment, solution was removed, and the scaffolds were rinsed with deionized water at room temperature for 24 h. Successively, they were dried in a vacuum desiccator and stored at room temperature for further modifications. Determination of Engrafted Amines. A ninhidryn-based procedure (Kaiser test, Aldrich) was employed to assess the amount of amino groups on the PCL-NH2 scaffolds. The samples were dissolved in kit solutions and heated at 100 °C for 15 min. Afterward, at room temperature methylenecloryde/ethanol solution was added to stabilize the blue compound and to bring the polymer mixture into solution. The absorbance was measured at 570 nm using an UV−vis spectrophotometer (Lambda 25, Perkin-Elmer). A calibration curve was obtained by 1,6-hexaneamine in methylenecloryde/ethanol solution. Each experiment was performed at least three times in triplicate. Peptide Conjugation. Peptide sequences were covalently grafted onto the surface in a two-step way by using an epoxy cross-linker in mild aqueous condition. First, the aminolyzed 3D scaffolds were treated with a 5% diethylene glycol diglycidyl ether (DGDGE) in a sodium carbonate solution (50 mM, pH 8.5) at room temperature gently shaking for 3 h. Subsequently, the tether solution was rinsed out, and the scaffolds were washed thoroughly with water. The conjugation step was performed adding 0.2 mg/mL of GRGDY (Inbios, Italy) in sodium carbonate solution (50 mM, pH = 8.5), and a scrambled peptide sequence (GYDGR) was used as a negative control for biological assessment. In both cases, an ethanolamine (2 mM) aqueous solution was used to deactivate unreacted groups. A scheme of the two-step procedure employed to graft GRGDY peptides on PCL fibers of the 3D scaffolds is reported in Figure 1. Determination of Conjugated Peptide. Micro-BCA assay (SigmaAldrich) was used to quantify the peptide density directly onto the bioactivated surfaces. The bicinchoninic acid assay is a biochemical

was precisely controlled through functional groups determination as well as chemical and physical parameter evaluation. Peptide surface density and distribution as well as penetration depth in the polymer substrate were deeply investigated through morphological and topological measurements. It was demonstrated that the conjugation of amine-terminated peptides by means of reductive amination after tether insertion may show a specific recognition of the solid signal to NIH3T3 integrin cell receptors suggesting a correct presentation of the peptide sequences.30 The possibility to extend this controlled two-step procedure to immobilize RGD motifs on 3D well-organized scaffolds could be very interesting, taking into consideration its effect on their macromechanical behavior. Accordingly, the present work mainly analyzes the effect of the surface treatment on the mechanical performance at different scales, also showing a morphological and analytical approach to study the functionalization/bioactivation treatment, the distribution of immobilized ligands, and the biological features of the 3D rapid prototyped scaffolds.

2. EXPERIMENTAL SECTION Materials and Methods. 3D Scaffolds Design and Preparation. 3D block-shaped scaffolds characterized by a length (l) of 7.0 mm, a width (w) of 7.0 mm, and a height (h0) of 7.8 mm, were fabricated through 3D fiber deposition technique, using a Bioplotter dispensing machine (Envisiontec GmbH, Germany) equipped with a CAD/CAM system. Poly(ε-caprolactone) (PCL) pellets (Aldrich, Mw = 65000) were initially placed in a stainless steel syringe and then heated at a temperature of 120 °C through a cartridge unit placed on the mobile arm of the XYZ plotter. As PCL reached the molten phase, a nitrogen pressure of 8−8.5 bar was applied to the syringe through a cap. 3D models were loaded on the Bioplotter CAD/CAM system. 3D scaffolds were obtained by alternatively extruding and depositing the polymer fibers with different angle steps between two successive layers, making two different patterns: 0°/90° and 0°/45°/90°/135°. The nozzle used to extrude PCL fibers was a stainless steel needle characterized by an inner diameter of 400 μm. Each scaffold was characterized not only by the fiber diameter (depending on the needle diameter and/or the deposition speed), but also by the fiber spacing (strand distance, i.e. center-to-center distance) and layer thickness, which influence the overall pore size. The values of strand distance were set to 640 μm, while the layer thickness was chosen to be 320 μm. A deposition speed of 50−55 mm/min was used. 3512

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assay for determining the total level of peptide immobilized on solid support in a solution using colorimetric techniques. This method combines the reduction of the Cu2+ to Cu1+ by peptide or protein in an alkaline medium with the highly sensitive and selective colorimetric detection of the Cu1+ using a BCA containing reagent. The purple colored reaction product of this assay is formed by the chelation51 of two molecules of BCA and one Cu1+. This water-soluble compound exhibits a strong absorbance at 562 nm that is linearly proportional to the peptide concentration. The amount of peptide bound was determined with a standard curve for known quantity of the same peptide. The number of peptide bonds and the presence of four amino acids (cysteine, cystine, tryptophan, and tyrosine) have been reported to be responsible for color formation in peptide samples when assayed with BCA. The kit solutions were prepared as described by the supplier in a reduced volume (1 mL) and added to the sample. After heating the mixture at 37 °C for 2 h, the absorbance was read at 562 nm and compared with a calibration curve, obtained each time by using a standard solution of GRGDY peptides in the range of concentration between 0.01 and 0.50 mM. Unmodified PCL 3D scaffolds were used as a negative control. Each experiment was performed at least three times in triplicate. Nanoindentation Tests. Nanoindentation tests were carried out on aminolyzed and not-aminolyzed PCL fibers, which were characterized by a diameter (D) of 340−360 μm and obtained through a Bioplotter dispensing machine (Envisiontec GmbH, Germany). All the tests were performed on dry specimens in a specific load range (1−5 mN), using a Nanotest Platform (Micromaterials, U.K.) with a diamond pyramidshaped Berkovich-type indenter tip. Trapezoidal load functions characterized by a loading−unloading rate of 300 μN/s and a peakload hold period of 20 s were imposed. Load-depth curves, values of hardness, and reduced modulus were evaluated. Hardness and reduced modulus were calculated using the methods introduced by Oliver and Pharr (1992) (see Supporting Information). In particular, hardness (H) was evaluated considering the applied peak load (Pmax) and the projected contact area (Ac) at the specified load, according to the equation

H=

The stress (σ) was evaluated as follows:

σ=

where F represents the force measured by the load cell divided and A0 (l·w) is the total area of the apparent cross section of the scaffold. The strain (ε) was defined as the ratio between the scaffold height variation Δh and its initial height h0: ε=

Pmax Ac

F A

where F is the force measured by the load cell and A (πD2/4) represents the fiber cross section. The engineering strain (ε) was evaluated as the ratio between the fiber elongation Δl and the initial fiber length l0 (i.e., the initial grips separation):

ε=

Δh h0

Spatial Distribution of Surface Treatment. Confocal laser scanning microscopy (CLSM) was used to investigate the penetration depth of treatment by LSM 510 Zeiss confocal inverted microscope equipped with a Zeiss 20X/3 NA objective and an argon laser. Each stage of bioactivation (aminolysis and peptide covalent coupling) were followed conjugating the surface samples with two different dyes. First, the aminolysis treatment were showed by coupling PCL-NH2 surfaces with 0.1 mg/mL of Rhodamine B isothiocyanate (RBITC, Sigma - R1755) in IPA overnight at 4 °C. Subsequently surfaces were rinsed thoroughly with IPA and then with water for 24 h to remove any non covalent bound dye molecule. In order to mimic the peptide route, Fluoresceinamine (FLUO, Fluka -07980) was linked to the epoxy-functionalized surfaces by first dissolving the dye in carbonate buffer at pH = 8.5 in the same conditions used for peptide conjugation. Finally, samples were rinsed with buffer followed by copious amounts of distilled water to remove any non-covalently linked molecules. The PCL fluorescent samples were then left to dry overnight in a vacuum desiccator before analysis. Unmodified PCL surfaces were equally processed as a control. Samples after RBITC or FLUO conjugation were respectively visualized using the characteristic wavelengths of RBITC (λex = 543 nm; λem = 572 nm) and FLUO (λex = 496 nm; λem= 518 nm). To compare the results, CLSM settings, in particular, laser power, pinhole aperture, detector gain, and amplifier offset, were kept constant for both kind of observations. The penetration depth of treatment was visualized by z stack acquisitions through the fiber by starting from the outer part. Intensity profiles of fluorescent dyes as a function of penetration depth were obtained along a line drawn in the radius direction (here indicated as the z-direction). Cell Adhesion Study. Mouse embryo fibroblasts NIH3T3 were maintained at 37 °C and 5% CO2 in Dulbecco’s modified Eagle medium (DMEM) supplemented with 10% fetal bovine serum (FBS, BioWhittaker, Walkersville, MD), 2 mM L-glutamine (Sigma, St. Louis, MO), 1000 U/l penicillin (Sigma, St. Louis, MO) and 100 mg/L streptomycin (Sigma, St. Louis, MO). In this study, 70−80% confluent cells were used. In particular, PCL, PCL-NH2, PCL-DGDGE-GYDGR, and PCL-DGDGE-GRGDY 3D scaffolds were sterilized with antibiotics and preincubated in serum-free medium for 16−18 h as reported in a previous work.30 After the incubation, 5 × 104 cells were seeded on all the different kinds of 3D scaffolds and grown in DMEM w/o FBS to avoid unspecific cell adhesion depending on serum protein adsorption. Scanning electron microscopy (SEM) was performed by a Leica 420 microscope in order to evaluate cell adhesion and shape. At different times after cell seeding, the 3D fiber-deposited scaffolds were rinsed with phosphate-buffered saline (PBS) and fixed with 2.5% glutaraldehyde (pH= 7.4) (Sigma-Aldrich, Italy) for 2h at room temperature. The cell-scaffold constructs were dehydrated in graded ethanol concentrations (from 50% to 100% v/v in ethanol), air-dried, gold sputtered and analyzed by SEM. Four different kinds of 3D scaffolds were studied: PCL, PCL-NH2, PCL-DGDGE-GYDGR, and PCL- DGDGE-GRGDY. Furthermore, the several cell-scaffold constructs were also analyzed (PCL, PCL-NH 2 , PCL-DGDGE-GYDGR, and PCL-DGDGEGRGDY) through CLSM. They were fixed with 4% paraformaldehyde for 20 min at room temperature, after 24 h from seeding, rinsed twice with PBS buffer, and incubated with 0.5% PBS-BSA (BSA = bovine serum albumin) to block unspecific binding. Actin microfilaments were

The projected contact area Ac is strongly related to the geometry of the indenter and it is calculated from the penetration depth. Tensile Tests. Tensile tests were performed on unmodified PCL and PCL-NH2 fibers according to the ASTM D3822 standard. The tests were performed on fibers (340−360 μm in diameter - D) in the standard atmosphere, which is 21 ± 1 °C and 65 ± 2% relative humidity. The gauge length was 20 mm and a rate of 48 mm/min was employed taking into consideration the table of the rate of extension. All the tests were carried out using an INSTRON 5566 testing machine. The engineering stress (σ) was calculated as follows: σ=

F A0

Δl l0

Compression tests. Compression tests were performed on the 3D fiber-deposited scaffolds in the form of block-shaped specimen characterized by a length (l) of 7.0 mm, a width (w) of 7.0 mm and a height (h0) of 7.8 mm. All the tests were carried out at a rate of 1 mm/min up to a strain value of 0.5 mm/mm, at 21 ± 1 °C and 65 ± 2% relative humidity, using an INSTRON 5566 testing system. 3513

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stained with phalloidin-tetramethylrhodamine B isothiocyanate (Sigma-Aldrich). Phalloidin was diluted in 0.5% PBS-BSA and incubated for 30 min at room temperature. The images were acquired by using a He−Ne excitation laser at the wavelength of 543 nm and a 20× objective. Alamar Blue assay (AbD Serotec Ltd., UK) was also performed on cell-scaffold constructs (PCL, PCL-NH2, PCL-DGDGE-GYDGR, and PCL-DGDGE-GRGDY) in order to evaluate cell viability. This method is based on a redox reaction that occurs in the mitochondria of the cells; the colored product is transported out of the cell and can be measured spectrophotometrically. The optical density was measured with a spectrophotometer (Sunrise; Tecan, Männedorf, Zurich, Switzerland) at wavelengths of 570 and 595 nm. The number of viable cells correlates with the magnitude of dye reduction and is expressed as a percentage of Alamar Blue reduction, according to the manufacturer’s protocol. Each experiment was performed at least three times in triplicate. Results. Micro-Computed Tomography. Micro-CT, which is an attractive single and nondestructive method to study the characteristics of scaffolds, has first allowed the morphological and architectural features of the 3D fiber-deposited PCL structures to be evaluated. This analysis has confirmed that well-organized PCL scaffolds have been obtained, showing precise pore size and shape, as well as a repeatable microstructure (Figure 2).

Figure 3. Amine and peptide density as a function of time of aminolysis treatment. Data are graphically reported as mean value, and bars represent the standard deviation. of the BCA reagent with tyrosine. Moreover, the BCA reagent requires at least a tripeptide in order to oxidize the peptide backbone. As in the case of amino groups, the density of peptides covalently coupled to the PCL has been calculated by taking into account the effective scaffold surface area. A peptide density of 20.13 ± 4.68 nmol/cm2 has been achieved (Figure 3). Spatial Distribution of Surface Treatment. To analyze the spatial distribution and penetration depth during each step of bioactivation, CLSM was employed. For this reason, two different dyes were coupled: RBITC was used to label free amino groups after aminolysis treatment, whereas in a second step, FLUO was conjugated to the epoxy-activated PCL surface to mimic the peptide behavior studying its distribution on the scaffold fiber surface. Figure 4 clearly shows the intensity profile of the fiber cross-section along the radius (here indicated as z-direction), evidencing a penetration depth of more than 140 μm for the aminolysis treatment. With regard to the FLUO bound surfaces, the total penetration depth of the treatment was about 80 μm, thus providing important information on the potential peptide penetration depth. The selected pore size and the fully interconnected pore network of the 3D rapid prototyped scaffold allowed us to optimize the mass transport properties (permeability and diffusion). The diffusion through the pores of the well organized structure is favored, thus obtaining a uniform distribution of the treatment within the inner structure. Accordingly, the amino groups grafted onto the fibers composing all the layers of the structure were uniformly distributed within the scaffolds (Figure 5) and would be used for peptide conjugation after tether (DGDGE) insertion. On the other hand, regarding the penetration depth of the single fibers, the differences in the central part of the fibers with respect to the crossing points of the fibers (fiber junctions) represent the only inhomogeneities. Nanoindentation Tests. Nanoindentation measurements on PCL fibers have displayed differences in terms of load-depth curves and, hence, of hardness values. Both aminolyzed (PCL-NH2) and unmodified (PCL) fibers have evidenced hardness (H) values that generally decrease as load increases from 1 to 5 mN. In particular, measurements on unmodified PCL fibers have evidenced hardness values spanning from 0.50 to 0.27 GPa in the load range investigated. These values are greater than those obtained for PCL fibers that were modified via aminolysis (0.1−0.03 GPa) (Figure 6a). This suggests that after aminolysis the fiber surface becomes softer. Consistently with hardness values, the reduced modulus (Er) of unmodified PCL fibers (4.2−1.2 GPa) is higher than that obtained from the PCL-NH2 ones (1−0.3 GPa) (Figure 6b).

Figure 2. 3D reconstructions obtained from Micro-CT analysis on PCL fiber-deposited scaffolds with 0°/45°/90°/135° (a) and 0°/90° (b) lay-down patterns. In particular, imaging analyses have evidenced a sufficient consistency between real and theoretical values, showing a mean fiber diameter of 340−360 μm and a center-to-center fiber distance of about 640−660 μm between two fibers in a common layer, as expected on the basis of the process/instrument parameters employed during the fabrication (i.e., needle inner diameter, deposition speed, fiber spacing...). Scaffold interconnectivity, which is normally defined as 100% × volume of interconnected pores/volume sum of interconnected and closed pores,52 has been also evaluated and found to be equal to 100% . Determination of Engrafted Amines and Conjugated Peptide. After aminolysis, an analytical determination of amino groups engrafted onto the surface of PCL fibers constituting the wellorganized scaffold was performed by a slight modification procedure based on the Kaiser test.30,53,54 With regard to the evolution of aminolysis treatment performed on the 3D scaffolds, a high amino-density of 161.3 ± 15.3 nmol/cm2 was reached after 30 min of aminolysis, and then a decrease of surface amines was observed over time (Figure 3). The amount of peptides immobilized on PCL fibers of the scaffolds was then evaluated using a one-pot colorimetric assay based on the BCA-Cu1+ purple color complex. This assay is widely employed to assess proteins both in solution and on adsorbing solid substrates in a very reproducible way and with high sensitivity (picomolar scale).51 The results obtained from micro-BCA assay have highlighted that the tyrosine present in our peptide provides a significantly different amount of color formation at 37 °C, thus suggesting a partial reaction 3514

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Figure 4. Intensity profiles of fluorescent molecules in the depth during the two-step conjugation procedure: (a) CLSM intensity profile as a function of depth in the z-direction of a PCL-NH-RBITC fiber; (b) CLSM intensity profile as a function of depth in the z-direction of a PCLDGDGE-FLUO fiber. unmodified ones, as can be also evidenced by the fluctuating stress values after the first local maximum of the stress−strain curve (Figure 7). Furthermore, even though the surface treatment strongly reduces the maximum strain, it seems to provide no differences in terms of modulus (E) and yield stress (σy) (Table 1). Compression Tests. Compression tests have evidenced that the aminolyzed 3D fiber-deposited scaffolds show stress−strain curves similar to those obtained for the unmodified ones. For both PCL-NH2 and PCL scaffolds, compressive stress−strain curves may be divided into three different regions: after an initial relatively stiff mechanical response, there is a region with lower stiffness, finally followed by another stiff portion, similar to the densification region of flexible foams (Figure 8). The compressive modulus (E) has been evaluated as the slope of the initial linear region of the stress−strain curve. To emphasize the effect of architecture and surface modification on the compressive mechanical performances of the 3D fiber-deposited scaffolds, values of compressive modulus and maximum stress are reported as mean value ± standard deviation for two different laydown patterns (0°/90° and 0°/45°/90°/135°), also before and after aminolysis (Table 2). Cell Adhesion Study. SEM and CLSM analyses have highlighted that NIH3T3 cells already adhere on the PCL fiber of the 3D scaffolds after 24 h from seeding, showing a morphology that drastically changes on the different samples (Figure 9). In particular, with regard to unmodified PCL, PCL-NH2, and PCLDGDGE-GYDGR scaffolds, cells appear more rounded, indicating a scarce adhesion to the substrate (Figure 9A/E, B/F, and C/G,

Figure 5. CLSM analysis: images of fibers belonging to two layers at different depths within a PCL-NH-RBITC scaffold characterized by a specific lay-down pattern (0°/90°). The reduced modulus may be considered a combined modulus obtained from nanoindentation tests as it is related to the Young’s moduli of both tip and sample, and to the their Poisson’s ratios. The tip properties are usually known, thus the Young’s modulus of a material can be evaluated from the reduced modulus if the Poisson’s ratio of the sample material is known (see Supporting Information). Tensile Tests. Tensile tests have shown a ductile behavior for both PCL-NH2 and unmodified PCL fibers. In particular, the stress−strain curves obtained show an initial linear region, then a little decrease in the slope occurs up to a local maximum stress value, followed by a decrease of the tensile stress. Then, a plateau-like region is evident, and finally a new increase of stress can be observed until failure is generally reached. During testing, the propagation of multiple necks were also observed along the aminolyzed fibers in comparison with the 3515

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Figure 6. Results obtained from nanoindentation tests on PCL and PCL-NH2 fibers: hardness (a) and reduced modulus (b) as a function of the applied load (1−5 mN). Data are graphically reported as mean value, and bars represent the standard deviation. The dashed lines are just a guide for the eye. bridging morphology was also frequent within the fully interconnected pore network. Furthermore, Figure 11 shows SEM micrographs of cells attached to fibers belonging to the inner layers of the PCL-DGDGE-GRGDY scaffolds after 4 days from seeding. Results from Alamar Blue assay suggest cell viability over time as the values reported in Table 3 increase with time. Discussions. The four steps initially carried out in our work can be summarized as follows: (a) design and fabrication of “morphologicallycontrolled” scaffolds through 3D fiber-deposition technique; (b) optimization of a one-step surface treatment, especially through the quantification of amino-groups and the analysis of the penetration depth; (c) conjugation and quantification of RGD peptides; (d) mechanical characterization performed to assess the effect on the surface and bulk PCL properties, as well as on the macromechanical compression performances of the structures. 3D rapid prototyped PCL scaffolds were fabricated through the 3D fiber-deposition technique, then functionalized via aminolysis reaction and bioactived by grafting RGD peptides to the surface-modified fibers. It has already been demonstrated that aminolysis may be considered an easy-to-perform chemical technique to engraft amino groups along polyesters chains.30 By treating the 3D well-organized PCL scaffolds with a 1,6-hexandiamine in an aprotic solvent at a 37 °C, a high density of amino groups were rapidly obtained onto the PCL fiber surface of the structure. The reaction starts by a nucleophilic attack onto the ester by an amino group at one end of diamine leading to the formation of an amide and leaving at the other end a free amino group emerging the PCL fiber surface. The functionalization pathway of the PCL scaffolds, which means the functionalization of the fiber surface of the porous structure, is reported in Figure 1. In addition to the new bond formation along the polymer surface, from the rupture of ester group were produced hydroxyl groups that could remain attached onto fiber surface or leached out during the washes. As summarized in Figure 1, the scaffold surface bioactivation was then performed by covalently grafting the GRGDY peptide via a homobifunctional cross-linker. DGDGE is a water-soluble epoxy crosslinker, that reacts quite well with amine in mild condition. The amount of peptides grafted to PCL fibers of the scaffolds was then quantified by using the one-pot colorimetric micro-BCA assay. The number of peptide bonds and the presence of four amino acids (cysteine, cystine, tryptophan and tyrosine) have been reported to be responsible for color formation in the peptide sample when assayed with BCA. Studies with tri- and tetra-peptides suggest that the extent of color formation is due to the presence of several functional groups.51 The advantages in using BCA methods include a compatibility with ionic and non ionic detergents, a stable working

Figure 7. Typical stress−strain curves obtained from tensile tests on PCL and PCL-NH2 fibers.

Table 1. Results from Tensile Tests Performed on PCL and PCL-NH2 Microfibers: Tensile Modulus (E), Yield Stress (σy), and Maximum Strain (εmax), Reported as Mean Value ± Standard Deviation Fibers

E (MPa)

σy (MPa)

εmax (mm/mm)

PCL PCL-NH2

570.5 ± 50.1 550.0 ± 48.6

25.0 ± 3.5 24.2 ± 3.7

12.7 ± 1.1 6.5 ± 0.5

respectively). Conversely, as for PCL bioactivated with RGD peptide, cells adhered and were well-spread on the fiber surface, evidencing a good interaction with the material (Figure 9 D/H). The effect of PCL functionalization/bioactivation in enhancing cell adhesion was further confirmed by actin cytoskeleton staining. This qualitative analysis indicated that cells better adhered on RGD bioactivated scaffolds compared to cells seeded on unmodified PCL, PCL-NH2, and PCL-DGDGE-GYDGR surfaces, where there is no evidence of a complete cytoskeleton organization. Just as an example, Figure 10 reports a SEM micrograph of a bridging event after 48 h from seeding (PCL-DGDGE-GRGDY). SEM analyses on the different 3D constructs allowed us to observe cells attached to single fibers or to multiple fibers, stretching across void spaces within the scaffold (bridging morphology). Even though cells attached to single fibers of the scaffold were mainly observed in all the analyzed 3D constructs, 3516

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Figure 8. Effect of lay-down pattern and surface modification via aminolysis on the mechanical properties of 3D rapid prototyped scaffolds: (a) Typical stress−strain curves for PCL scaffolds characterized by two different lay-down patterns (0°/90° and 0°/45°/90°/135°), before (PCL) and after aminolysis (PCL-NH2); (b) Stress−strain curves reported up to a strain level of 0.04 mm/mm in order to better highlight the different initial stiffness of the 3D morphologically controlled structures.

Table 2. Effect of Lay-down Pattern and Surface Modification via Aminolysis on the Mechanical Properties of 3D Rapid Prototyped Scaffoldsa compressive modulus E (MPa) lay-down pattern 0°/45°/90°/ 135° 0°/90°

maximum stress σ (MPa)

PCL

PCL-NH2

PCL

PCL-NH2

63.0 ± 4.7

61.1 ± 5.1

12.3 ± 1.1

12.0 ± 1.3

89.1 ± 6.9

87.9 ± 8.1

13.5 ± 1.3

13.2 ± 1.5

Compressive modulus and maximum stress reported as mean value ± standard deviation, for PCL scaffolds characterized by two different lay-down patterns (0°/90° and 0°/45°/90°/135°), before (PCL) and after aminolysis (PCL-NH2). a

Figure 10. Cell adhesion study after 48 h from seeding: SEM micrograph (PCL-DGDGE-GRGDY), bar 50 μm. determine the amount of the functional groups. The quantification of the different groups should be calculated from a standard curve of an appropriate substance. In particular, the one-step colorimetric method has been considered to quantify the surface concentration of GRGDY peptides covalently

reagent and a tolerance to the presence of compounds that could interfere. As described by Tyllianakis et al.,55 this method can be used to determine the total solid supports functionalized with cysteine and tyrosine, requires only one incubation step, and allows one to

Figure 9. Cell adhesion study after 24 h from seeding: SEM micrographs (A: PCL; B: PCL-NH2; C: PCL-DGDGE-GYDGR; D: PCL-DGDGEGRGDY), bar 20 μm; CLSM images of phalloidin staining of microfilaments (E: PCL; F: PCL-NH2; G: PCL- DGDGE-GYDGR; H: PCL-DGDGEGRGDY). 3517

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Figure 11. Cell adhesion study after 4 days from seeding: different SEM micrographs (PCL-DGDGE-GRGDY). Wojtowicz et al. (2010)58 implanted GFOGER-coated wellorganized PCL 3D scaffolds into critically sized femoral defects in rats. If compared to uncoated PCL scaffolds, and empty defects, GFOGER-coated scaffolds seemed to enhance osteoblastic differentiation and subsequent bone repair without the aid of implanted cells, genetic material or growth factors. Taubenberger et al. (2010)59 showed that the exposure of RGD-motifs in collagen may present a mechanism to induce tissue remodeling and wound healing. In particular, partially denatured collagen I guided the binding of a variety of integrins, thereby improving cell adhesion and subsequent cellular migration and spreading. However, it is worth noting that none of the above-mentioned studies has assessed the effect of the surface modification on the mechanical behavior of the 3D rapid prototyped scaffolds. Consequently, trying to fill this gap present in the literature, nanoindentation, tensile, and compression tests were carried out with the aim to highlight the effect of aminolysis on the surface and bulk material properties, as well as on the macro-mechanical performances of the 3D fiber-deposited PCL scaffolds. Advanced materials as well as biological tissues show hierarchical structures with particular features down to the nanometer or micrometer scale. For this reason, a technique that can probe mechanical properties at these scales has to be considered. In this context, nanoindentation is emerging as a valuable mechanical testing technique for biomaterials. Hardness and microhardness testing (e.g., Vickers and Knoop indentation)60,61 have been already considered to investigate the mechanical properties of hard tissues such as teeth and bones.62−65 However, nanoindentation enhances upon the spatial, force, and displacement resolutions of these traditional techniques, thus providing a powerful tool to study tissues and biomaterials with submicrometer resolution. Nanoindentation is also useful for measuring mechanical properties of microstructural features within bulk samples, characterizing the properties of individual constituents within composite or heterogeneous samples, or mapping mechanical properties across a sample surface. Because of its small probe size, nanoindentation can be used to measure local material properties in small, thin, and heterogeneous samples. Nanoindentation, which is an instrumented or depth-sensing indentation, involves, for example, the application of a controlled load to the surface inducing local surface deformation. Load and displacement are monitored during the loading and unloading phases. Thus, properties such as hardness and reduced modulus are calculated from the unloading curves using wellestablished equations. Considering its typical working force range and displacement range (1 μN−500 mN and 1 nm−20 μm, respectively),60,66 nanoindentation technique surely bridges the gap between atomic force microscopy (AFM) and macroscale mechanical testing. Accordingly, we decided to carry out nanoindentation measurements on PCL fibers in order to highlight the differences in terms of hardness and modulus values. Results from nanoindentation tests have evidenced that after aminolysis the fiber surface becomes softer, since the hardness values decrease if compared to those obtained for the unmodified PCL (Figure 6a). This softening effect could be ascribed to the reduction of the entanglement density due to the surface treatment. It is well-known that hardness and modulus of a polymer depend on its structure,

Table 3. Alamar Blue Assay Performed on Cell-Scaffold Constructs (PCL, PCL-NH2, PCL-DGDGE-GYDGR, and PCL-DGDGE-GRGDY): Results (Mean Value ± Standard Deviation) Reported at 1 and 5 Days Alamar Blue reduction (%) sample

1 day

PCL PCL-NH2 PCL-DGDGE-GRGDY PCL-DGDGE-GYDGR

8.2 ± 1.3 9.7 ± 2.9 10.5 ± 1.8 10.3 ± 1.6

5 days 18.3 22.1 27.9 20.7

± ± ± ±

4.9 5.4 5.2 3.5

bound on the fibers of the 3D PCL scaffolds. The quantification was carried out by a standard curve using a known concentration of 1,6hexanediamine. In our work, a peptide density of 20.13 ± 4.68 nmol/ cm2 has been achieved, while a value of 2.81 ± 0.35 nmol/cm2 has been assessed by Causa et al. (2010) in their study.30 As reported in the literature, a minimum RGD density of 1.0 × 10−15 mol/cm2, that corresponds to a spacing of 140 nm between peptide ligands, seems to be sufficient to promote cell spreading, while a density of 1.0 × 10−14 mol/cm2 has been found to favor the formation of focal contacts.30,38 The value of peptide density obtained in our work would seem well beyond the minimum value reported in the literature. Anyway, it is difficult to numerically compare our results obtained for 3D rapid prototyped scaffolds with those reported in literature for polymeric sheets. At each step of functionalization/bioactivation, the spatial distribution and penetration depth was properly evaluated through CLSM considering two different dyes (RBITC and FLUO). In the literature, many works have already shown how several surface modifications and RGD immobilization may improve the wettability and/or the biological performances of PCL scaffolds. For example, Mattanavee et al.56 performed the immobilization of various biomolecules (i.e., collagen, chitosan, and GRGDS peptide) on the surface of electrospun PCL fibrous scaffold, using N,N′-disuccinimidylcarbonate as the coupling agent. They demonstrated that the aminolyzed and biomolecule-immobilized PCL fibrous scaffolds became more hydrophilic as evidenced by dynamic water contact angle measurements. Furthermore, cell culture experiments with different cell lines have highlighted that among the bioactivated electrospun PCL scaffolds, the type I collagen- and GRGDSimmobilized scaffolds evidenced the greatest ability to promote cell attachment proliferation.56 Furthermore, Zhang et al.57 previously extended a simple method to immobilize RGD peptide on 2D PCL films to the case of 3D wellorganized scaffolds, investigating the bone marrow stromal cell (BMSC) behavior. Their modification strategy allowed one to obtain a successful Arg−Gly−Asp−Cys (RGDC) immobilization on 3D rapid prototyped PCL scaffolds via aminolysis and a heterobifunctional cross-linker sulfosuccinimidyl 4-(N-maleimidomethyl)cyclohexane-1carboxylate (sulfo-SMCC). In particular, Zhang et al.57 have widely studied BMSC attachment, cellular distribution, signal transduction, and survival on their RGD-modified PCL scaffolds, demonstrating that the modification elicits specific cellular responses and improves cell− biomaterial interactions. 3518

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molecular weight, and number of segments between entanglements.67 In particular, values of reduced modulus and hardness are related to polymer chain flexibility (i.e., physical and chemical entanglements). The rigidity of amorphous region of the polymers depends on the density of entanglements among molecular chains, which represent the topological restriction of molecular motion by other chains. The indenter tip mainly interacts with molecular chains of the surface, which are relatively more flexible than those in the bulk because of the greater mobility of surface chains. Within the deformation zone under the tip, the superposition of simultaneous responses of several molecular chains determines the nanoindentation behavior. It is wellknown that in the first stage, the aminolysis starts preferentially at the amorphous regions of the polymer causing the scission of the chains present on the surface, thus reducing the density of entanglements. Also taking into account the penetration depth of the treatment, as a consequence of the scission of the chains undergoing functionalization, the flexibility and mobility of surface molecular chains increases, and the surface of the aminolyzed fiber becomes softer. On the other hand, tensile tests performed on PCL fibers have basically evidenced how the surface treatment strongly reduces the maximum strain without altering the values of modulus and yield stress (Figure 7 and Table 1). Although nanoindentation and tensile measurements on PCL fibers have shown some differences, results from compression tests have suggested that fixing a specific lay-down pattern the aminolyzed 3D fiber-deposited scaffolds and the unmodified ones show similar mechanical properties. Furthermore, their stress−strain curves (Figure 8) are similar to those of flexible foams. Anyway, in contrast to the typical behavior of a flexible foam,68 the central zone is not a plateau, i.e., it does not have zero slope but just a lower one if compared with the other two portions of the stress−strain curve. This behavior is also consistent with that already reported for 3D fiber-deposited PCL scaffolds.24,69 As expected, the architecture, which means the specific lay-down pattern used (0°/90° or 0°/45°/90°/135°), also influences the mechanical behavior of the 3D PCL scaffolds in compression. The lay-down pattern strongly affects the mechanical behavior of the 3D fiber-deposited PCL scaffolds, especially in terms of initial stiffness. As reported in Table 2, before the surface modification via aminolysis, PCL scaffolds characterized by a 0°/90° pattern exhibit a compressive modulus (89.1 ± 6.9 MPa), which is greater than that obtained for a 0°/45°/90°/135° pattern (63.0 ± 4.7 MPa). At a strain value of 50%, a maximum stress of 13.5 ± 1.3 MPa and 12.3 ± 1.1 MPa has been evaluated for 0°/90° and 0°/45°/90°/135° patterns, respectively. Furthermore, the surface treatment via aminolysis does not negatively affect the macromechanical behavior of the 3D fiberdeposited scaffolds, as evaluated through compression tests. For example, after aminolysis, PCL scaffolds characterized by a 0°/90° pattern have shown values of compressive modulus (87.9 ± 8.1 MPa) that are similar to those achieved before the surface treatment (89.1 ± 6.9 MPa), as well as a maximum stress of 13.2 ± 1.2 MPa compared to 13.5 ± 1.3 MPa obtained for the corresponding not- aminolyzed structures. Similar observations might be made for the 3D fiber-deposited PCL scaffolds with a 0°/45°/90°/135° pattern, taking into account the results numerically reported in terms of modulus and maximum stress, before and after the surface treatment via aminolysis (Table 2). Consequently, the surface treatment via aminolysis reduces the hardness, but it does not negatively affect the compressive mechanical behavior of the rapid prototyped scaffolds. Unlike non-porous structure, this effect may be ascribed to the specific load transfer mechanism occurring in the 3D fiber-deposited scaffolds, where fibers act as load-bearing beams. The main contribution to the macromechanical performance is basically due to the crossing points of the fibers (fiber junctions) undergoing compression (i.e., a “column-like behavior”) which seem to be less affected by the treatment. Although the above-mentioned steps are crucial, the knowledge of cell-material interactions turns out to be a key element in designing 3D advanced multifunctional scaffolds with suitable morphology and

properties that are able to guide cell adhesion. For this reason, as a final step of this research, in order to analyze the effect of functionalization/bioactivation at the cell-material interface, the interaction of 3D fiber-deposited scaffolds with fibroblast cells was studied through SEM and CLSM. Results from these analyses have qualitatively allowed to show the cytoskeleton organization and to demonstrate that cells better adhered on RGD bioactivated scaffolds.

3. CONCLUSIONS RGD motifs are being widely considered to design biomimetic surfaces that could trigger a specific function in cell behavior at the cell-material interface. Accordingly, cell adhesion should be suitably enhanced and tailored since it represents the basic feature in the cell-material interaction. Previous works have already evidenced that aminolysis represents an easy route to introduce primary amines with high yield that can be easily optimized. Here, the controlled two-step procedure proposed by Causa et al.30 for 2D PCL films was extended to immobilize RGD motifs on 3D rapid prototyped scaffolds, designing 3D advanced scaffolds. Unlike other works on 3D RGD-modified scaffolds, the present study basically focuses on the effect of the surface modification on the mechanical properties at different scales, however, showing an approach to analyze the functionalization/bioactivation, the distribution of immobilized ligands, and the biological features. The determination of amines and peptides effectively engrafted on the fiber surface of 3D scaffolds and the treatment penetration depth were suitably assessed and discussed. Nanoindentation and tensile measurements carried out on PCL fibers allowed us to underline the effect of the functionalization on the surface and bulk properties. More importantly, the surface modification did not negatively affect the macromechanical behavior of the 3D rapid prototyped scaffolds as evaluated through compression tests. On the other hand, results from cell adhesion study evidenced that the conjugation of peptides through aminolysis and tether insertion enhanced NIH3T3 cell adhesion and spreading.

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ASSOCIATED CONTENT

S Supporting Information *

Nanoindentation: theory and methods, adaptation for indentation of polymers. This material is available free of charge via the Internet at http://pubs.acs.org

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AUTHOR INFORMATION

Corresponding Author

*Mailing address: Institute of Composite and Biomedical Materials, National Research Council, P.le Tecchio 80, 80125, Naples, Italy. Telephone number: +39-081-2425925; fax number: +39-081-2425932; e-mail address: [email protected]. Author Contributions §

These authors contributed equally to this work.

Notes

The authors declare no competing financial interest.

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ACKNOWLEDGMENTS The authors wish to thank Mr. Mario De Angioletti and Mr. Rodolfo Morra for performing nanoidentation measurements and mechanical tests, respectively.

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