5 Additive Manufacturing of Wet-spun Polymeric Scaffolds for Bone Tissue Engineering

Chapter V

5 Additive Manufacturing of Wet-spun Polymeric Scaffolds for Bone Tissue Engineering

Dario Puppi

, Carlos Mota

, Matteo Gazzarri

, Dinuccio Dinucci

, Mairam Myrzabekova

, Antonio Gloria

, Luigi Ambrosio

and Federica Chiellini

1

Laboratory of Bioactive Polymeric Materials for Biomedical and Environmental Applications (BIOLab), via Vecchia Livornese 1291, 56010 San Piero a Grado (Pi). Department of Chemistry and Industrial Chemistry, University of Pisa, Italy.

2

Institute of Composite and Biomedical Materials, National Research Council, Naples, Italy.

Abstract

An Additive Manufacturing technique for the fabrication of three-dimensional polymeric scaffolds, based on wet-spinning of poly(ε-caprolactone) (PCL) or PCL/hydroxyapatite (HA) solutions, was developed. The processing conditions to fabricate scaffolds by a layer-by-layer approach were optimised by studying their influence on fibres morphology and alignment. Two different scaffold architectures were designed and fabricated by tuning inter-fibre distance and fibres staggering. The developed scaffolds showed good reproducibility of the internal architecture characterized by highly porous, aligned fibres with an average diameter in the range 200 – 250 µm. Mechanical characterization showed that the architecture and HA loading influenced the scaffold compressive modulus and strength. Cell culture experiments employing MC3T3-E1 preosteoblast cell line showed good cell adhesion, proliferation, alkaline phosphatase activity and bone mineralization on the developed scaffolds.

Keywords: tissue engineering, scaffolds, wet-spinning, additive manufacturing, polycaprolactone.

5.1 Introduction

Bone tissue engineering (TE) is one of the most promising approaches to be used as alternative to

the conventional autogenic or allogenic surgical techniques for bone tissue repair. Scaffold-based

TE strategies involve the use of a biodegradable, porous scaffold that serves as structural template

to fill the tissue lesion and to support cell-cell interactions and extracellular matrix (ECM)

formation. Under optimal conditions, cells harvested from donor tissues, including adult or stem cells, are expanded in culture and associated with a scaffold of synthetic and/or natural origin. The scaffold/cells construct is then implanted in the targeted site where the defect can be regenerated as consequence of a good interaction with the host tissue.[Puppi et al., 2010]

Macro and micro-structural properties of the scaffold affect not only cells survival, signalling, growth, propagation and reorganization, but play also a major role in modelling cell shape and gene expressions, both related to cell growth and preservation of native phenotypes [Karageorgiou and Kaplan, 2005; Leong et al., 2003]. Since the first pioneering experiments carried out by Langer and Vacanti more than 20 years ago [Langer and Vacanti, 1993; Vacanti et al., 1988], several studies have reported different materials processing techniques for the fabrication of polymeric scaffolds with a macro- and micro-architecture suitable for TE applications. These include, among others, solvent casting combined with particulate leaching, freeze drying, gas foaming, melt moulding, fibre bonding, phase separation, electrospinning and rapid prototyping techniques [Puppi et al., 2010].

Wet-spinning is a non-solvent induced phase inversion technique allowing for the production of a continuous micrometric polymer fibre through an immersion precipitation process: a polymeric solution is injected directly into a coagulation bath containing a poor solvent for the polymer, and the solution filament solidifies because of polymer desolvation caused by solvent/non-solvent exchange [Puppi et al., 2011b]. Among other techniques for manufacturing polymeric fibres employed in biomedical applications, wet-spinning has been mostly used to process natural polymers, such as chitin and chitosan [Tuzlakoglu and Reis, 2008], which cannot be formed by other spinning techniques. A growing body of literature has recently proposed wet-spun microfibres for TE applications, including chitosan fibres [Tuzlakoglu et al., 2004], braided poly(L-lactic acid) (PLLA)/chitosan fibres [Zhang et al., 2007], starch-based non-woven fibrous meshes [Leonor et al., 2011; Pashkuleva et al., 2010; Tuzlakoglu et al., 2010], poly(ε-caprolactone) (PCL) fibres [Williamson and Coombes, 2004], star poly(ε-caprolactone) (*PCL) non-woven fibrous meshes [Puppi et al., 2011a; Puppi et al., 2011b]. In particular, assemblies of wet-spun fibres, obtained by either physical bonding of prefabricated fibres or by a single-step method involving the continuous, randomly-oriented deposition of the solidifying fibre in the coagulation bath, have been shown to possess a three-dimensional (3D) structure with high and interconnected porosity suitable for TE purposes [Leonor et al., 2011; Pashkuleva et al., 2010; Puppi et al., 2011a;

Puppi et al., 2011b; Tuzlakoglu et al., 2004; Tuzlakoglu et al., 2010]. However, these fabrication

methods suffer from lack of structure reproducibility as well as control over external shape and

internal morphology [Puppi et al., 2011a; Puppi et al., 2011b].

Additive manufacturing (AM), which can be defined as “the process of joining materials to make objects from 3D model data, usually layer upon layer” [Keriquel et al., 2010], has been extensively applied for the fabrication of 3D scaffolds by employing different techniques, such as stereolitography, selective laser sintering, 3D printing and fused deposition modeling (FDM) [Woodruff and Hutmacher, 2010]. Thanks to their ability to produce porous polymeric matrices with reproducible and customized microstructure and macroshape, such techniques represent a significant breakthrough in scaffolds manufacturing. In particular, over the past decade and since the first work reported by Hutmacher on scaffolds fabricated by FDM [Hutmacher, 2000], a number of studies have been published on melt extrusion-based AM techniques for application in TE [Domingos et al., 2009; Mota et al., 2011; Wang et al., 2004; Woodfield et al., 2004]. These techniques involve the fabrication of layers of parallel strands with different orientation one on top of the other, by depositing with a predefined pattern an extruded filament of a polymer melt.

PCL is a semicristalline polymer that has been widely investigated in bone tissue regeneration applications because of its biocompatibility and slow degradation [Koh et al., 2006; Rai et al., 2007; Williams et al., 2005; Woodruff and Hutmacher, 2010]. However, the implantation of PCL substitutes into bone defects typically can present various drawbacks, such as lack of integration with the surrounding tissue because of an inflammatory reaction, encapsulation into fibrous tissue and mechanical strength reduction associated with material degradation [Kokubo et al., 2003;

Schiller and Epple, 2003; Taylor et al., 1994]. The incorporation of hydroxyapatite (HA), a synthetic calcium phosphate ceramic that mimics the natural apatite composition of bones and teeth, into biodegradable polyesters has been investigated as an effective means of improving the osteoconductivity and mechanical properties of bone implants and creating a pH buffer against the acidic degradation products of the polymeric matrices [Kikuchi et al., 2004; Kim et al., 2004; Koh et al., 2006; Ural et al., 2000; Wutticharoenmongkol et al., 2007].

Aim of the present work was the development of an AM technique allowing enhanced controlled

over internal and external architecture of microfibrous polymeric scaffolds fabricated by wet-

spinning. By exploiting a computer-assisted wet-spinning system, the processing conditions for the

fabrication of 3D scaffolds, with different architectures and made of either PCL or PCL/HA

composite, were optimised. The developed scaffolds were characterized for their morphology and

elemental composition by means of scanning electron microscopy (SEM), under backscattered

electron imaging and microanalysis, and micro-computed tomography (µCT), as well as for their

mechanical compression properties using a uniaxial testing machine. In vitro cell culture

experiments employing MC3T3-E1 murine preosteoblast cells were carried out in order to evaluate

the scaffolds cytocompatibility. Cell response, in terms of viability, proliferation, morphology,

differentiation and bone mineralization, was investigated by tetrazolium salts (WST-1 cell proliferation reagent), confocal laser scanning microscopy (CLSM), alkaline phosphates activity (ALP) and alizarin red staining (ARS) respectively.

5.2 Materials and Methods

5.2.1 Additive manufacturing of wet-spun scaffolds

5.2.1.1 Materials

Poly(ε-caprolactone) (PCL, CAPA 6800, Mw = 80000 g·mol

) was kindly supplied by Solvay (Italy), hydroxyapatite (HA) nanoparticles (size < 200 nm) were purchased from Sigma-Aldrich (Italy). All the solvents and chemical reagents were purchased from Sigma-Aldrich (Italy) and used as received.

5.2.1.2 Preparation of polymeric solutions

PCL was dissolved in acetone at 35 °C, under gentle magnetic stirring for 3 h to obtain homogeneous solutions of various concentrations. For the production of composite scaffolds, the desired amount of HA was added to the polymer solution and left under vigorous stirring for 2 h to achieve a homogeneous dispersion. On the base of some preliminary investigations, the weight ratio between HA and PCL was chosen to be 25%.

5.2.1.3 Fabrication of 3D polymeric scaffolds

The polymeric solution was placed in a glass syringe connected to a blunt tip stainless steel needle (Gauge 23) through a plastic tube. By using a programmable syringe pump (KDS100, KD Scientific, MA) the solution was injected at a controlled feeding rate directly into a bath of ethanol.

The X-Y movement of the needle and the Z movement of the build platform were controlled by an

in-house made computer-controlled system allowing for the production of 3D structures layer-by-

layer (Figure 1).

Figure 1 - a) Scheme of computer-aided wet-spinning apparatus; b) layer-by-layer process.

Various processing parameters, such as polymer concentration (C

), solution feed rate (F), initial distance between the needle tip and the bottom of the beaker (Z

), inter-fibre needle translation distance (d

), X-Y translation velocity (V

) and Z interlayer translation distance (d

), were investigated in order to optimize the process for the fabrication of scaffolds with different internal architectures. Rectangular prism-shaped PCL scaffolds with the base measuring 15 × 15 mm and the height measuring around 5 mm were produced for morphological and compressive mechanical characterization. Wet-spun scaffolds were removed from the coagulation bath after production, then placed in a vacuum chamber for 48 h and finally stored in a desiccator.

5.2.2 Morphological characterization

5.2.2.1 Scanning electron microscopy (SEM)

Samples were cut from the produced scaffolds and characterized by scanning electron microscopy (SEM, model JEOL LSM5600LV, Japan) under backscattered electron imaging and elemental microanalysis. The average fibre diameter and inter-fibre distance were determined by means of ImageJ 1.43u software on micrographs with a 50X magnification; morphological parameters were calculated over 30 measurements per specimen, taken from randomly selected fields. Scaffold elemental composition was analysed on five random areas (80 × 50 µm) of each sample.

5.2.2.2 Micro-computer tomography (µCT)

Micro-computed tomography (µCT) was carried out on scaffolds using a SkyScan 1072 system

(Aartselaar, Belgium). A rotational step of 0.9° over an angle of 180° was imposed. Cross-sections

and 3D models of PCL and PCL/HA structures were reconstructed using SkyScan’s software package, Image J software, Rapidform and Materialise Mimics.

5.2.3 Mechanical characterization

Compression tests were carried on scaffolds in order to assess their mechanical behaviour. Block- shaped specimens were characterized by a length (l) of 15.0 mm, a width (w) of 15.0 mm and a height (h

) of about 5.0 mm. Six samples for each kind of scaffold were characterized at a rate of 1 mm·min

and up to a strain value of 0.5 mm·mm

, using an INSTRON 5566 testing machine. The stress σ was defined as the measured force (F) divided by the total area of the apparent cross- section of the scaffold (A

= l·w):

A

= F

σ ₍₁₎

whilst the strain ε was evaluated as the ratio between the structure height variation ∆h and its initial height h

:

h 0

h

= ∆

ε (2)

5.2.4 Biological characterization

5.2.4.1 Cell seeding

Scaffold samples were placed in a 24 wells plate, sterilized under UV light for half an hour each side and then washed with 70% ethanol:water solution for 3 h. After ethanol removal, scaffolds were extensively washed with phosphate buffer saline (PBS 1X, pH 7.4), containing a penicillin/streptomycin solution (1%), and then left overnight at 37 °C in a 5% CO

. The solution was then substituted with complete culture medium and samples were incubated for 24 h before cell seeding. Mouse calvaria-derived preosteoblastic cells MC3T3-E1 (subclone 4 ATCC CRL- 2593) were obtained from the American Type Culture Collection (ATCC CRL-2593) and cultured as monolayers in Alpha Minimum Essential Medium (α-MEM, Sigma), containing ribonucleosides, deoxyribonucleosides and sodium bicarbonate, supplemented with L-glutamine (2 mM), fetal bovine serum (10%), penicillin:streptomycin solution (100 U/ml:100µg/ml) (1%) and antimycotic.

The cultures were maintained at 37 °C and 5% CO

. Confluent cells at passage 25 were trypsinized

(0.25% trypsin containing 1 mM EDTA), centrifuged and resuspended in complete medium.

Subsequently 0.5 x 10

cells per scaffold were seeded onto scaffolds in a 24 well plate and, after one hour of incubation at 37°C and 5% CO

, 600 µl of complete medium were added to each well, followed by incubation in a humidified atmosphere at 37°C. In order to induce and promote cells osteoblastic phenotype expression, some scaffolds were cultured in osteogenic medium obtained by supplementing the α-MEM with ascorbic acid (0.3 mM) [Quarles et al., 1992]. The medium was replaced every 48 h.

5.2.4.2 Cell viability and proliferation

Cell viability and proliferation were measured by using the cell proliferation reagent WST-1 (Roche) after 7, 14, 21, 28, 35 and 43 d of cell culturing. The test is based on the mitochondrial enzymatic conversion of the tetrazolium salt WST-1 into formazan, the soluble product. The assay was performed by incubating cell-seeded scaffolds for 4 h with the WST-1 reagent, diluted 1:10, at 37°C and 5% CO

. Measurements of formazan dye absorbance were carried out with a microplate reader (Biorad) at 450 nm, with the reference wavelength at 655 nm.

5.2.4.3 Alkaline phosphatase (ALP) activity

ALP activity was determined in cultured MC3T3-E1-scaffold constructs after 7, 14, 21 and 28 d of

cell culturing. The measurement was assessed with a colorimetric method based on the conversion

of p-nitrophenyl phosphate into p-nitrophenol by the ALP enzymatic activity. Scaffolds were

washed three times with PBS and then placed into 1 ml of a lysis buffer, containing Triton X-100

(0.2%), magnesium chloride (5 mM) and trizma base (10 mM) at pH 10. Samples underwent

freezing-thawing cycles by keeping at –20°C and subsequently at room temperature (RT)

[Wutticharoenmongkol et al., 2007]. This process was repeated three times in order to extract the

intracellular ALP [Zhou et al., 2010]. Afterwards, a volume of 20 µl of supernatant was taken from

the samples and added into 100 µl of p-nitrophenyl phosphate substrate (Sigma). A standard

calibration, prepared dissolving ALP from bovine kidney (Sigma) in the same lysis buffer, was

added to the substrate and the reaction was left to take place at 37°C for 30 min. The reaction was

stopped by adding 50 µl of 2 M NaOH solution and after 5 min absorbance was measured at 405

nm in a microplate reader. The molar concentration of ALP was normalized with the total protein

content of each sample, which was measured using Bradford protein assay (Pierce). The amount of

the proteins was calculated against a standard curve. The results for ALP assay were reported as

nano-moles (nmol) of converted substrate/(mg of protein·minute)

.

5.2.4.4 Mineralized matrix deposition analysis by alizarin red staining (ARS)

The mineralized matrix deposition was analysed by using the ARS method [Ozkan et al., 2009;

Stanford et al., 1995] after 7, 14, 21, 28, 35 and 43 d of cell culturing. Scaffolds fixed with 3.8% p- formaldehyde at RT for 30 min, were stained with 2% Alizarin Red solution pH 4 for 10 min. After the dye incubation, samples were extensively rinsed with sterile de-mineralized, in order to remove the dye excess. To quantify the amount of calcium deposited on the scaffolds, the red matrix precipitate was dissolved in 10% cetylpyridinium chloride solution pH 7, and the optical density of the solution was read with a micro-plate reader at 565 nm. A standard calibration of ARS was carried out in cetylpyridinium chloride buffer. Unseeded scaffolds were treated with the same procedure, as blank.

5.2.4.5 Cell morphology investigation by confocal laser scanning microscopy (CLSM)

Morphology of MC3T3-E1 cells grown on the prepared scaffolds and 3D culture organization were investigated by means of CLSM. Cells were fixed with 3.8% p-formaldehyde for 30 min in PBS 1X, permeabilized with a PBS 1X/Triton X-100 solution (0.2%) for 15 min and incubated with a solution of 4’-6-diamidino-2-phenylindole (DAPI) (Invitrogen) and phalloidin-AlexaFluor488 (Invitrogen) in PBS for 45 min at room temperature in the dark. After dye incubation, samples were extensively washed with PBS and observed by including specimen in-between two glass cover slips. All steps of the above procedure were performed under gently shaking on an orbital shaker in order to enhance solution penetration into scaffold structure. A Nikon Eclipse TE2000 inverted microscope equipped with an EZ-C1 confocal laser and Differential Interference Contrast apparatus was used to analyse the samples (Nikon). A 405 nm laser diode and an Argon Ion Laser (488 nm emission) were used to excite respectively DAPI and FITC fluorophores. Images were captured with Nikon EZ-C1 software with identical instrumental settings for each sample. Images were further processed with The GIMP (GNU Free Software Foundation) image manipulation software and merged with Nikon ACT-2U software.

5.2.5 Statistical analysis

All the in vitro biological tests were performed on triplicate samples for each material. Quantitative

data were presented as mean ± standard deviation (SD). Data sets were screened by one-way

ANOVA and a Tukey test was used for post hoc analysis; significance was defined at p < 0.05.

5.3 Results and Discussion

Techniques based on a layered manufacturing strategy represent an effective approach for the control at the microscale over the internal architecture, external shape and size of TE scaffolds. In this study, an innovative AM technique for the fabrication of wet-spun polymeric scaffolds was developed. The processing conditions for the manufacturing of 3D wet-spun PCL and PCL/HA composite scaffolds with different internal architectures were investigated, and the developed scaffolds were characterized for their morphology, compressive mechanical properties and cytocompatibility using the MC3T3-E1 murine preosteoblast cell line.

5.3.1 Additive manufacturing of 3D wet-spun structures

The scaffold fabrication process involved the extrusion of a polymeric solution through a X-Y translating needle that was immersed into a coagulation bath (Figure 1). A layer composed of parallel fibres was fabricated by depositing the solidifying solution filament with a predefined pattern, and 3D architectures were built up with a layer-by-layer process by fabricating layers with different fibre orientation (0-90°lay-down pattern) one on top of the other.

By optimizing the processing parameters, two different scaffold architectures were developed (Figure 2): one (PCL

) was obtained with an X-Y inter-fibre needle translation distance (d

) of 1 mm and 0.5 mm staggered fibre spacing between successive layers with the same fibre orientation;

the other one (PCL

) was obtained with a d

of 0.5 mm.

Figure 2 - a) PCL

(left) and PCL

(right) scaffold architectures, b) 3D PCL scaffolds fabricated by AM

(64 layers).

Besides d

, the most influent processing parameters on fibres alignment and morphology were polymer concentration (C

), F, starting needle tip-collection plane distance (Z

) and deposition velocity (V

) (Table 1). PCL/HA composite scaffolds with the above described two architectures were successfully fabricated by processing a suspension of HA nanoparticles in PCL solution and applying the processing conditions optimised for plain PCL scaffolds.

Table 1 - Processing conditions, fibre diameter and inter-fibre distance of the different kinds of developed scaffold. Morphological parameters are expressed as average ± standard deviation.

Scaffold C

[w/v]

F [ml·h

]

Z

[mm]

V

[mm·min

]

Fibre diameter [µm]

Inter-fibre distance [µm]

PCL

10% 1.4 5 300 202.1 ± 11.7 872.5 ± 65.4

PCL

8% 1 1 170 238.4 ± 13.4 256.3 ± 36.4

PCL/HA

10% 1.4 5 300 240.7 ± 13.4 807.5 ± 41.8 PCL/HA

8% 1 1 170 241.7 ± 21.5 262.3 ± 59.5

Due to its superior rheological and viscoelastic properties with respect to various resorbable-

polymer counterparts, PCL has been widely investigated during the past two decades for the

manufacturing of a large range of scaffold structures, including nanofibre meshes, foams, knitted

textiles and rapid prototyped constructs. As recently reviewed by Woodruff and Hutmacher

[Woodruff and Hutmacher, 2010], a vast array of AM techniques with a layered manufacturing

strategy has been proposed for the fabrication of PCL scaffolds, such as those based on laser and

UV light sources (i.e. stereolithography, selective laser sintering, solid ground curing), 3D printing,

melt extrusion-based techniques (i.e. FDM and precision extruding deposition) and direct writing

techniques. In addition, wet-spinning of PCL solutions was recently proposed for the fabrication of

3D non-woven microfibrous scaffolds through a single-step process [Puppi et al., 2011a; Puppi et

al., 2011b]. However, this technique, besides requiring a continuous handiwork, suffers from poor

control over scaffold microstructure and shape. The AM technique developed during the present

activity allows for overcoming these disadvantages by exploiting a computer-assisted wet-spinning

system that enables to design and manufacture by a layer-by-layer approach, PCL-based 3D

constructs with improved control over scaffold microarchitecture and shape. This technique does

not require high temperatures and involves the use of solvents classified with low toxic level, thus

allowing for the loading of bioactive agents without compromising their bioactivity.

5.3.2 Morphological analysis

SEM analysis using backscattering electron imaging showed good reproducibility of internal architecture and good degree of fibres alignment for the two kinds of PCL architecture developed.

The fibres presented a highly porous morphology both in the outer surface and in the cross-section, with a pore size of few micrometres (Figure 3). Moreover, PCL/HA composite scaffolds revealed a morphology close to that of plain PCL scaffolds.

Figure 3 - Representative backscatter SEM micrographs of the two scaffold architectures: top view and cross-

section of PCL

(top) and PCL

(bottom) scaffolds. Insert high magnification micrographs show

porosity of the outer surface and cross-section of the fibres.

As reported in Table1, PCL

scaffolds showed significantly smaller fibre diameter (202.1 ± 11.7 µm) in comparison with PCL

scaffolds (238.4 ± 13.4 µm); HA loading resulted in larger diameter in the case of PCL/HA

(240.7 ± 13.4 µm) while it did not affect significantly dimension in the case of PCL/HA

(241.7 ± 21.5 µm). The inter-fibre distance (defined as the minimum distance between two adjacent fibres within the same layer) was in the range 700 – 950 µm for PCL

and in the range 200 – 350 µm for PCL

.

Figures 4 shows representative images obtained from µCT analysis of PCL/HA scaffolds with the two developed architectures. This analysis highlighted the well-defined morphology and the architectural features of the structures. In particular, it showed the good degree of fibre alignment in the inner part of the scaffolds (Figures 4a and 4b) and confirmed that the developed AM technique is a powerful tool to manufacture scaffolds characterized by a repeatable structure.

Figure 4 - Representative images obtained from µCT analysis: cross-section of a) PCL/HA

and b) PCL/HA

; 3D reconstructions of c) PCL/HA

and d) PCL/HA

.

Figure 5 shows representative electron back scattering microanalysis spectra of PCL/HA composite

scaffolds. Chemical composition analysis of fibre surface revealed the presence and quite

homogeneous distribution of phosphorus and calcium elements, unequivocally associated with the

presence of HA. However, some white spots (Figure 5b) were observed in high magnification

micrographs of fibre surface. The marked differences in intensity of analogue energy peaks between the spectra of wide fibre surface areas (Figure 5a) and those of white spots (Figure 5b) revealed that they were mainly composed of HA elements .

Figure 5 - Representative EBS micrograph and elemental analysis of PCL/HA scaffold: (a) spectra of the wide fibre surface area; (b) spectra of a white spot onto fibre surface.

Upon immersion of a homogeneous polymeric solution into a coagulation bath, typically a dense, non-porous layer (skin) is formed at the interface with the non-solvent because of instantaneous non-solvent diffusion into the polymer solution [Tsay and McHugh, 1992; Wienk et al., 1996]. The spongy morphology of the outer surface and cross-section of the fibres constituting the scaffolds developed during the present work is likely due, as suggested by previous studies on phase inversion mechanism [Barton et al., 1997; Wienk et al., 1996], to a delayed liquid–liquid demixing.

In comparison with dense strands fabricated by melt-based AM techniques (e.g. FDM), such highly

porous fibre morphology can present some advantages in influencing, in addition to the

biodegradation rate and the mass transfer associated to drug release phenomena, the mechanisms

regulating cell adhesion and proliferation [Karageorgiou and Kaplan, 2005]. The HA

microaggregates detected on the fibres surface are likely due to the not perfectly homogenous

suspension of HA nanoparticles in the PCL solution. Future research will investigate whether such

aggregates were leached out during the pre-treatment stages before in vitro studies or in vivo

implantation (e.g. sterilization, scaffold preconditioning).

5.3.3 Mechanical characterization

The effect of HA loading and scaffold architecture on compressive mechanical properties of the developed scaffolds was tested using a uniaxial testing machine (strain rate = 1 mm·min

, maximum strain = 0.5 mm·mm

). As shown in Figure 6, the stress-strain curves of the manufactured PCL and PCL/HA scaffolds were characterized by an initial linear region at low values of strain, followed by two further regions with different stiffness.

Figure 6 - a) Typical stress-strain curves obtained for scaffolds compressed at a rate of 1 mm·min

up to a

strain value of 0.5 mm·mm

: a) PCL

structure; b) PCL

structure.

Data reported in Table 2 showed that PCL

displayed significantly higher compressive modulus and maximum stress in comparison with PCL

structure, as well as PCL/HA composite scaffolds in comparison with plain PCL scaffolds.

Table 2 - Compressive modulus and maximum stress, reported as mean value ± standard deviation, for the different kinds of scaffolds developed. The samples were compressed at a rate of 1 mm·mm

1 up to a strain value of 0.5 mm·mm

.

Scaffold Compressive Modulus [MPa]

Maximum stress (50% strain) [MPa]

PCL

0.12 ± 0.05 0.26 ± 0.05

PCL/HA

0.21 ± 0.05 0.39 ± 0.03

PCL

0.60 ± 0.20 0.34 ± 0.10

PCL/HA

0.90 ± 0.24 0.47± 0.11

The compressive mechanical properties of the developed scaffolds are quite different from those of load-bearing bone tissues which experience high stresses and low strains during in vivo physiological loading. The stiffness of this kind of scaffolds would need thus to be significantly improved in order to broaden their applications in bone regeneration. Their mechanical behaviour is consistent with that reported in previous works on melt extrusion-based AM of PCL scaffolds [Bartolo et al., 2011; Gloria et al., 2009; Hutmacher et al., 2001; Kyriakidou et al., 2008; M.

Domingos et al., 2009], showing stress-strain curves characterized by three regions with different slope, although they differ in the absence of a central plateau region with roughly constant stress as well as in the much lower values of compressive modulus and mechanical strength. These differences might be related to the highly porous fibre morphology that can cause a different mechanical answer of the single fibre. The enhanced mechanical properties as consequence of HA loading corroborate the results of a number of studies on polymer/inorganic composite scaffolds [Rezwan et al., 2006]. The higher compressive modulus and strength of PCL

scaffolds in comparison with those of PCL

can be explained with their higher fibre packing density. Future studies will be dedicated to the tuning of internal architecture parameters (i.e. inter-fibre distance, fibre staggering, lay-down pattern) and to assess the effect of phase inversion conditions (e.g.

solvent/non-solvent system, temperature) on fibre morphology and consequently on the overall

scaffold mechanical performance.

5.3.4 Biological characterization

As shown in Figure 7, the developed scaffolds were able to support the proliferation of MC3T3-E1 cells in the two investigated typologies of scaffolds (plain PCL and PCL/HA) and growing conditions (non-osteogenic and osteogenic). Despite the low values of cell proliferation at day 7, an increasing trend was evident for all the tested samples, with an average peak of proliferation between 28 and 35 d of culture. Cells grown onto plain PCL

scaffolds reached slight higher values of proliferation if compared with the ones grown onto plain PCL

, while PCL/HA scaffolds showed slight lower cell proliferation in comparison with unloaded scaffolds. The two different culture conditions instead did not show a marked difference in cell proliferation.

Figure 7 - Proliferation of MC3T3-E1 cell line cultured on PCL scaffolds by AM: (a) plain PCL constructs;

(b) PCL/HA constructs. non ost: non osteogenic medium; ost: osteogenic medium.

Cell-material interactions and cell seeding density play a crucial role on cell attachment, thus influencing cell proliferation in the first week of culture [Kommareddy et al., 2010]. The observed low cell proliferation at day 7 of culture was probably due to the large pore size of the prepared scaffolds that was not effective in retaining a significant number of cells during the seeding procedure. However, despite the low initial number of attached cells, the investigated scaffolds were able to support the proliferation of the MC3T3-E1 cell line during the following weeks. The decrease in cell viability at day 43 was likely due to the limited available surface left on the construct for the expansion of the culture, as corroborated by CLSM analysis. The slower cell proliferation observed on HA-loaded constructs correlates with the detected amount of ALP suggesting a more pronounced differentiation process on HA-loaded scaffolds [Quarles et al., 1992].

Results showed comparable levels of ALP activity for both structures (plain PCL and PCL/HA) and geometries (PCL

and PCL

), with a time-dependent increasing trend (Figure 8). As expected the investigated constructs cultured in osteogenic medium showed higher levels of ALP [Alcaín and Burón, 1994; Zhou et al., 2010]. The presence of HA in the PCL scaffolds could have helped MC3T3-E1 to trigger expression of the enzyme. This phenomenon was markedly evident at day 28 in the scaffolds with the 0.5 mm geometry, demonstrating the efficacy of the synergy between the HA and the osteogenic medium. [Calvert et al., 2005; Wutticharoenmongkol et al., 2007].

The ALP values detected from the MC3T3-E1 cells cultured on both plain and HA loaded PCL

structures suggested that the marked starting point of the differentiation process took place between

14 and 21 d of culture. These time-dependent changes in ALP production indicated the division of

osteoblast development into two distinct stages [Quarles et al., 1992]. The initial phase was

characterized by active replication of undifferentiated cells. In this regard, during the first two-three

weeks, cultures displayed a rapid increase in cell number (Figure 7), but these immature cells

expressed low levels of ALP (Figure 8) and failed to mineralize [Hoemann et al., 2009]. The

second phase was characterized by a diminished cell proliferation, and expression of bone cell

phenotype. The down-regulation of the replication, quite evident for all the samples and the culture

conditions in the period between 21 and 28 d of culture, was coupled to the expression of high

levels of ALP, a marker of mature osteoblast function. Indeed, ALP activity, low during active

replication, increased significantly with the onset of growth arrest [Quarles et al., 1992]. The

obtained results supported the analysis of the formation of the mineralized ECM that increased at

day 21, in coincidence with the increasing values of ALP. In fact, ALP is believed to be involved in

the hydrolysis of pyrophosphate (inhibitor of the mineralization process) towards inorganic

phosphate, that induces the mineralization occurring by means of apatite formation [Beck et al., 1998; Calvert et al., 2005; Wutticharoenmongkol et al., 2007].

Moreover, the delay in the maximum expression of ALP can be explained on the basis of the cellular three-dimensional organization. In fact, it is likely that the up-regulation for the production of ALP was triggered by cellular contacts and/or expression of ample amounts of early matrix proteins as type I collagen, fibronectin and transforming growth factor beta (TGF-β1). In these 3D structures cells need in the “initial phase” to form multi-layered clusters reaching the adequate confluence that acts on the decrease of the rate of proliferation and on the expression of the bone isoform of ALP [Park, 2010; Wutticharoenmongkol et al., 2007].

Figure 8 - ALP activity detected on MC3T3-E1 cells cultured on (a) plain PCL and (b) PCL/HA scaffolds.

non ost: non osteogenic medium; ost: osteogenic medium.

During the first three weeks of cell culturing appreciable values of matrix mineralization were not observed, confirming data reported by the literature that consider the ECM mineralization a late stage indicator of osteoblastic phenotype [Whited et al., 2011]. Since day 21, cells cultured on scaffolds showed to be involved in the production of high levels of mineralized ECM, with an increasing trend during the culture time (Figure 9).

Figure 9 - Calcium deposits on MC3T3-E1 cells cultured on (a) plain PCL and (b) PCL/HA scaffolds

investigated by mean of ARS method. non ost: non osteogenic medium; ost: osteogenic medium.

The values of ARS detected from plain PCL

constructs treated with ascorbic acid were particularly remarkable. Moreover, cells grown on PCL/HA scaffolds displayed levels of ARS roughly 10 times higher in respect to the amount detected in cultured plain PCL samples.

These findings were consistent with other studies in which polymer-HA composite scaffolds increased bone ECM formation and mineralization of preosteoblasts when compared to scaffolds without a biomimetic apatite component [Whited et al., 2011].

Fluorescent staining of cytoskeleton and nuclei showed morphology of cells grown on cultured scaffolds. After 7 d of cell culture, microscopic observation showed diffused presence of preosteoblast cells adherent on fibres surface with a variable shape and spreading. F-actin organization was consistent with early stages of cell adaptation to the material [Hutmacher et al., 2001], exhibiting great stress fibres stretched along the cytoplasm, and a low cell number coherent with the quantitative proliferation data (Figure 10a). Yet cellular presence was higher on lower layers fibres likely due to cells slid down after the seeding procedure. No apparent difference was detectable for number and morphology of cells by comparing the two structures (PCL

and PCL

) and the two different materials (PCL and PCL/HA) at early stages. The analysis of samples cultured for longer times showed a progressive increase in cells colonizing the polymeric structure in all the experimental conditions. By the third week, cells started to exhibit a polygonal morphology and to form discrete groups to yield, at the fourth week, large cell clusters extensively covering fibres surface and spanning through the layers with complex inter-cellular connections (Figures 10b and 10c).

These observations were in accordance with the differentiation pathway proposed for the

preosteoblasts in vitro, after an early growing latency, morpho-functional cellular aggregates are

developed and single cell morphology is not distinguishable [Quarles et al., 1992]. At the fifth

week, cultured samples exhibited a nearly full cellular colonization of available fibre surface by a

wide continuous cell culture net. In many cases single cells, or multiple cell structures, were

observed to bridge adjacent fibres layers (Figures 10d and 10f). The formation of a complex

multicellular coverage did not allow the observation of peculiar morphological differences in cell

morphology in cultured scaffolds in any of the experimental conditions. In PCL

scaffold

samples (Figure 10d), due to the shorter inter-fibre distance, cellular covering of the gap between

fibres was achieved earlier than in PCL

scaffolds (Figure 10e). Moreover, no further differences

were observed in 35 d cultured PCL/HA scaffolds compared to the analogous of plain PCL (Figure

10f).

Figure 10 - CLSM microphotographs resuming MC3T3-E1 cell cultured on PCL and PCL/HA based

scaffolds, at different weeks end-point. PCL

scaffold samples cultured for 7 days (a), 21 days (b), and 28

days (c) show increasing cellular population. PCL

(d) and PCL/HA

(e) comparison indicates no

difference at day 35 in cellular presence. PCL

at 35 days (f) exhibits a lower inter-fibre cell bridging. Scale

in (a) applicable to all picture sets.

5.4 Conclusions

The main result attained during the present activity was the development of a layer-by-layer AM technique allowing for the production by computer assisted wet-spinning of customized tissue engineered PCL or PCL/HA composite scaffolds. In particular, it was shown that it is possible to manufacture 3D structures with controlled and reproducible internal microarchitecture and external shape by collecting with a predefined pattern a solidifying filament of PCL solution into a coagulation bath. Tuning of manufacturing parameters enabled to customize the internal architecture features, such as inter-fibre distance and fibres staggering. Differently to what commonly observed in scaffold produced by melt extrusion-based AM, the fibres constituting the scaffolds showed a highly porous morphology, due to the phase inversion process, that worsened the scaffold mechanical performance. However, it is reasonable to hypothesize that spongy fibre morphology could positively affect mass transfer related to the exchange of nutrients and bioactive agents. The developed PCL and PCL/HA scaffolds with the two investigated structures were able to support the adhesion and proliferation of MC3T3-E1 preosteoblast cells that colonized the inner parts of structures during 43 d cell culturing experiments. The tested scaffolds were able to support the mechanism of differentiation of preosteoblast cells stimulating the production of high levels of mineralized ECM. In addition, HA loaded into the scaffolds and ascorbic acid addition to culture medium significantly enhanced bone mineralization and ALP activity.

The present study opens new possibilities for the fabrication of 3D structures with a layered manufacturing approach by employing other biodegradable polyesters that are well suited for wet- spinning processing (e.g. *PCL [Puppi et al., 2011a; Puppi et al., 2011b], PLLA [Gao et al., 2007]

and poly(lactic-co-glycolic acid) [Mack et al., 2009]). In addition, an ongoing work that will be published in a forthcoming paper is also addressed to the development by computer-assisted wet- spinning of anatomically-shaped, clinically-sized scaffolds for the in vivo treatment of bone tissue defects.

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