Chapter IX
9 Development of Novel Scaffolds for Otologic Surgery Applications
Carlos Mota1,2, Serena Danti3, Clemens Blitterswijk2, Stefano Berrettini3, Federica Chiellini1, Emo Chiellini1, Lorenzo Moroni2
1 Laboratory of Bioactive Polymeric Materials for Biomedical and Environmental Applications (BIOLab), Department of Chemistry and Industrial Chemistry, University of Pisa, via Vecchia Livornese 1291, 56010 San Piero a Grado (Pi), Italy.
2 Department of Tissue Regeneration, MIRA Institute for Biomedical Technology and Technical Medicine, University of Twente, 7500 AE Enschede, PO BOX 217, The Netherlands
3 Laboratory of Otologic Tissue Engineering & Temporal Bone Dissection (OTOLab), Department of Neurosciences, University of Pisa, Via Paradisa 2, 56124 Pisa, Italy
Abstract
In the present work, additive manufacturing and electrospinning techniques were used to produce partial ossicular replacement prosthesis (PORP), posterior canal wall and tympanic membrane (TM) scaffolds. For the production of the scaffolds, a biocompatible and biodegradable block copolymer - poly(ethylene oxide terephthalate) and poly(butylene terephthalate - was used.
Scaffolds morphology was assessed by means of scanning electron microscopy. PORP surface topography was examined by means of atomic force microscopy. A tissue engineering approach, combining PORP scaffolds and human mesenchymal stem cells (hMSCs), was investigated performing in vitro cell culture experiments. Qualitative analyses were performed to evaluate cell distribution on the scaffolds. Quantitative analyses showed that scaffold were able to retain an average of 40% of the initially seeded cells and the cell number remained constant along the cell culture experiment. Moreover, a higher osteogenic differentiation of hMSCs was observed when osteogenic culture medium was used. These preliminary results showed the potentialities of additive manufacturing and electrospinning techniques to produce middle ear scaffolds. Moreover, biological assessment showed the capability of the PORP scaffolds to support adhesion, proliferation and differentiation of hMSCs towards an osteoblastic phenotype.
Keywords: middle ear, partial ossicular replacement prosthesis (PORP), tympanic membrane (TM), posterior canal wall (PCW), tissue engineering, polyActive, human mesenchymal stem cells (hMSCs), 3D fibre deposition, electrospinning.
9.1 Introduction
The human ear is anatomically divided into three sections: the outer ear (containing the auricle and external auditory meatus), the middle ear (comprising the tympanic membrane (TM) and the ossicular chain) and the inner ear (containing the cochlea and the vestibular apparatus). The middle ear function is to convey sound pressures, canalised through the external auditory meatus, via the TM and ossicular chain into the inner ear fluids [Haberman, 2004]. Conductive hearing loss generally induces reduction or impairment of the auditory function at middle ear level representing one of the major hearing diseases. This problem can occur due to TM perforations or to the loss of function of the ossicular chain. Causes such as, infection, trauma, birth defects and degenerative diseases are associated to the conductive hearing loss [Danti et al., 2010]. Treatment of this disease generally involves surgical techniques known as myringoplasty, performed in cases of TM perforations, and tympanoplasty, performed in cases of ossicular damages. In the latter option, TM has usually to be reconstructed. Ossiculoplasty is the part of a tympanoplasty intervention specifically devoted to ossicle replacement. In severe cases (e.g. totally damaged middle ear), surgeons perform a total reconstruction using several types of prostheses. The development of a complete middle ear prosthesis set, including a canal wall prosthesis, a TM/malleus prosthesis, and an incus stapes prosthesis was performed by Grote et al [Grote, 1984;
Grote, 1985; Grote, 1995]. Clinical trials with these prosthesis were performed in 15 patients [Grote, 1995] and satisfactory results were observed on 12 patients after one year follow-up, however, after three years, obliteration and infection were observed in seven cases. The long-term failure was attributed to extensive middle ear diseases observed before the initial surgery.
The failure rates vary depending on the surgical schools, the biomaterials (grafts of synthetic biomaterials), and are generally associated to persistent infections, cholesteatoma or poor fit of the prosthesis [van Blitterswijk et al., 1990].
Pioneer work performed by Wullstein in the early 1950s [Wullstein, 1956] introduced alloplastic artificial material (vinyl-acryl plastic) for ossicular chain reconstruction. Other alloplastic materials, such as polymers, metals, ceramics and combinations of these, were also proposed for ossicular chain prosthetization [Danti et al., 2010; Hildmann et al., 2006]. Despite of the good results obtained with the alloplastic materials, extrusion, resorption and dislocation of the prosthesis may occur.
A novel approach proposed recently by Danti et al, suggests a tissue engineering (TE) strategy for the development of a biohybid prosthesis combining scaffolds with human mesenchymal stem cells (hMSCs) [Danti et al., 2010; Danti et al., 2009]. The TE approach will allow better results in terms of biostability, functionality and biointegration of the scaffolds because of the presence of a living cellular phase previously incorporated in vitro. One of the most promising approaches in TE relies on the utilization of patient's own cells that are collected by means of a small biopsy, expanded and seeded into a three-dimensional (3D) scaffold and re-implanted into the patient [Ikada, 2006;
Moroni et al., 2008b]. HMSCs collected from bone marrow have been drawing the attention of the scientific community because of the possibility of being expanded in culture while preserving their multipotency. Moreover, when hMSCs are cultured under proper conditions, they differentiate into end-stage cell types, such as osteoblasts, chondrocytes, myocytes, tenocytes, adipocytes, fibroblasts, and other connective tissue cells [Caplan, 2007]. Furthermore, hMSCs secrete a broad spectrum of bioactive macromolecules that have immunoregulatory function and create structured microenvironments in the damaged area [Caplan, 2007].
3D scaffolds are also an important pillar of the TE approach. Several are the techniques, generally divided into conventional and Additive Manufacturing (AM) techniques, to produce scaffolds for TE application. AM techniques are a group of techniques that allow the production of scaffolds with controlled 3D external geometry and internal architecture of pores. Fused Deposition Modeling (FDM) is one of the AM techniques commonly used for the production of scaffolds for hard tissue replacement, because of the possibility of producing these types of structures with customized mechanical properties [Moroni et al., 2006b; Zein et al., 2002]. Moreover, it is possible to produce biomimetic custom-made scaffolds from computer-aided design (CAD) models obtained from the patients by magnetic resonance image (MRI) or computer tomography (CT) techniques. Three-dimensional fibre deposition (3DF) has been extensively studied for the production of scaffolds for diverse TE applications (e.g. bone, cartilage, osteochondral defects, neotrachea) [Agarwal et al., 2009; Moroni et al., 2007; Moroni et al., 2005, 2006a; Moroni et al., 2008c]. This AM technique follows a similar principle of FDM allowing the melt processing of a wider range of polymers.
Electrospinning (ES) is another attractive and simple technique to produce non-woven fibrous scaffolds proposed TE applications [Moroni et al., 2008a]. The similarity of the electrospun fibres with the natural extracellular matrix (ECM) fibrils makes this technique ideal to produce ECM substitutes [Moroni et al., 2008a]. The high surface area to volume ratio is an advantage of electrospun meshes that provides a large amount of biding sites for cells membrane receptors due to the nanoscaled architectures [Stevens and George, 2005].
Polyactive™ is a block copolymer composed of poly(ethylene oxide terephthalate) and poly(butylene terephthalate) (PEOT/PBT) with a wide range of physico-chemical and mechanical properties that can be tailored by changing the copolymer composition [Deschamps et al., 2004;
Moroni et al., 2005]. This copolymer has been extensively studied in the last decade for TE and drug-delivery applications, because when varied the molecular weight of the starting poly(ethylene glycol) segments and/or the weight ratio of poly(ethylene oxide terephthalate) (PEOT) and poly(butylene terephthalate) (PBT) blocks, it is possible to tailor properties such as wettability, swelling, biodegradation rate, protein adsorption, and mechanical properties [Moroni et al., 2005].
PEOT/PBT copolymers have Food and Drug Administration approval and CE mark for diverse medical devices. PEOT/PBT was firstly proposed for the production of scaffolds for the TM in the early 1990s [Grote et al., 1991]. In vivo studies on a rat model showed that PEOT/PBT films were covered with an epidermis and epithelium layers after two and four weeks and after one year the polymer degradation was approximately 54%. Moreover, clinical trials performed in 15 patients [Grote, 1995] showed good results in 12 cases after one year follow-up. However, after three years, obliteration and infection were observed in seven patients. The long-term failure was attributed to the extensive middle ear diseases observed before surgery [Grote, 1995].
The aim of the present study was to produce middle ear scaffolds for otologic surgery. Three types of scaffolds, partial ossicular replacement prosthesis (PORP), TM and posterior canal wall (PCW), were developed and produced by means of 3DF and ES techniques after the optimization of the processing parameters. PEOT/PBT was used to prepare all the developed scaffolds. TM scaffolds with a controlled pattern produced by 3DF were coated with a thin electrospun mesh. Moreover, the collection of continuous meshes was performed using a rotating drum. The characterization of the electrospun meshes and 3D scaffold morphology was performed by means of scanning electron microscopy (SEM). Furthermore, PORP surface topography was characterized by means of atomic force microscopy (AFM). PORP scaffolds were biological characterized and cultured with hMSCs during 21 days. Quantitative (DNA assay and alkaline phosphatase (ALP) activity assay) and qualitative analysis (methylene blue, SEM, and histological analysis) were performed on the cultured PORP scaffolds at different time-points.
9.2 Materials and Methods 9.2.1 Materials
PolyActive™ was provided by PolyVation BV (The Netherlands). The commercial designation of this random block copolymer follows an aPEOTbPBTc nomenclature, where a represents the
molecular weight (MW, g·mol-1) of the poly(ethylene oxide) (PEO), b and c represent the weight ratios of the poly(ethylene oxide terephthalate) (PEOT) and poly(butylene terephthalate) (PBT) respectively. The block copolymer used in this study was 300PEOT55PBT45. Chloroform (CHCl3) (Merck KGaA, Germany) and hexafluorisopropanol (HFIP) (Biosolve BV, Valkenswaard, The Netherlands) were used as received.
9.2.2 Scaffolds design
9.2.2.1 Partial ossicular replacement prosthesis (PORP) scaffold
For the manufacturing of PORP-like scaffolds PEOT/PBT copolymers was used. The geometry selected for these structures followed the 3D model reported earlier [Berrettini et al., 2011; Danti et al., 2010]. The model with a mushroom-like geometry was selected according to the commercial synthetic hydroxyapatite PORP model currently used in otologic surgery. The head surface of the PORP scaffold is placed in contact with the TM while the other extremity is placed in contact with the stapes (or stirrup) bone. In this study, the production of PORP scaffolds with a coaxial head with respect to the stem and with three different sizes (1:1; 1.5:1 and 2:1 ratios) was investigated.
The 1:1 scale scaffold was designed with the same dimensions of commercial hydroxyapatite PORP (L = 3.5mm; Dh = 4 mm; H = 0.8 mm; Ds = 2.2 mm). The 3D models of the PORP scaffolds were designed using Rhinoceros® software (McNeel, Seattle, USA) (Figure 1) and manufactured.
PORP scaffolds were produced using a 3DF technique.
Figure 1 – 3D models of PORP scaffolds modelled by means of Rhinoceros® software with three different sizes (1:1; 1.5:1 and 2:1 ratios). Main PORP dimensions are reported in the image.
9.2.2.2 Posterior canal wall (PCW) scaffold
PCW (Figure 2) was modelled according to a commercial available prosthesis. The 3D modelling of the PCW was performed implementing the following steps: a) accurate measuring of the features and geometry of a commercial sample using a digital calliper and micrometre (Mitutoyo Corp., Japan) and reference structures (planar ruler, precision set square), and b) generation in SolidWorks software (Dassault Systemes, France) of a 3D CAD model. For the generation of the 3D model, a portion of hollow ellipsoid matching the overall size of the sample was generated and further steps were performed: a) deformation of the primitive to match the geometry of the sample; b) 3D cutting of concave features as in the sample, and c) filleting of edges with appropriate curvature radii. Two cycles of iteration of the above mentioned steps were made, in order to verify the exactness of the 3D model with respect to the physic sample. The achieved accuracy was less than 0.1 mm. The overall dimensions of the CAD model were 10x20x20 mm. For the production of the PCW scaffolds, PEOT/PBT material was used and processed with 3DF technique.
Figure 2 – 3D CAD model designed from a commercial PCW prosthesis by means of SolidWorks software.
Image contains overall PCW dimensions.
9.2.2.3 Tympanic membrane (TM) scaffolds
For the production of the TM scaffold a two-step approach was used, combining the 3DF and ES techniques (Figure 4). A first step comprised the manufacturing of a tympanic membrane pattern (TMp) structure by means of 3DF technique (see subchapter 9.3.2.1). A second step was the coating of the TMp with a thin electrospun mesh (see subchapter 9.3.2.2). Both elements that constituted the TM scaffolds (TMp and electrospun mesh) were produced using PEOT/PBT copolymer. Additional details of the processing parameters are described in the scaffold fabrication subchapter. Three experimental designs were developed for the production of TM substitutes. A pattern with 15 mm of diameter was developed with the aim of the validation of static cell culture experiments (Figure 3d, 11a-c). Moreover, a pattern with a diameter of 25 mm was developed for dynamic cell culture experiments (Figure 11d-f). Flat electrospun meshes without TMp were also produced.
Figure 3 – Schematic representation of the different collagen fibre that compose the human TM: a) Radial fibres, b) Circular fibres and c) Parabolic fibres and d) simplified TMp designed with Rhinoceros composed of radial fibres with contour and circular fibres. Image a-c) adapted from Lim [Lim, 1995].
Figure 4 – Two-step strategy implemented for the production of TM scaffolds: a) first step comprised the production of TMp by means of 3DF technique, and b) subsequent coating of the TMp with a thin electrospun mesh. Image adapted from Moroni et al [Moroni et al., 2008c].
9.2.3 Scaffold fabrication
For the fabrication of PORP, PCW and TMp scaffolds, 3DF technique was used. For the manufacturing of TM scaffolds, the prepared TMp were coated with an electrospun mesh by means of ES techniques. These techniques and the processing parameters used for the production of the different scaffolds are described below.
9.2.3.1 Three-dimensional fibre deposition (3DF) technique
The equipment used to produce the PORP, PCW and TMp scaffolds was 3D-Bioplotter™
(envisionTEC GmbH, Germany). This AM equipment uses the principle of Fused Deposition Modeling (FDM) assisted by nitrogen (N2) gas pressure to promote the extrusion of a molten polymer. The parameters that influence the production of the 3D scaffolds are: temperature, N2
pressure, deposition velocity and extrusion nozzle diameter. A detailed description of the system was previously reported by Moroni et al [Moroni et al., 2006a; Moroni et al., 2008c]. Briefly, the system is composed of a stainless steel syringe were the PEOT/PBT was loaded and heated at a temperature of 205ºC. After obtaining the complete polymer melt, a N2 pressure of 5,5 bar was applied while an electrovalve controlled the extrusion of the polymer. For the fabrication of the PORP and PCW scaffolds, two nozzles with different internal diameters (I.D.) were used (G25, I.D.=250 µm and G27, I.D.=200 µm). The deposition velocity was optimised and varied from 56 to 196 mm·min-1. Each 3D model of the PORP and PCW scaffolds was loaded into the equipment PrimCAM software (Primus Data, Switzerland) and the deposition patterns for each scaffold were calculated. The fibre spacing (d2), defined as the distance between successive fibres in the same layer, was varied from 300 µm to 600 µm and the layer thickness (d3) was set to 150 µm. The fibre diameter (d1) obtained according to the nozzle diameter used. Scaffolds were built layer upon layer with a 0-90º architecture, where fibres were deposited with a 90º step between successive layers.
The PORP scaffolds will hereafter be referred using the following nomenclature “xPORPy”, where
“x” refers to the scale size (1; 1.5 and 2) and “y” to the value of d2 in microns.
For the fabrication of the TMp a deposition velocity of 56 mm·min-1 was used. The different layer arrangements (Figure 3d, 11a,e) that composed the TMp were designed using Rhinoceros® software (McNeel, Seattle, USA) and exported to the equipment. The TMp was composed of two layers (circular and radial fibres) and the layer thickness (d3) was set to 150 µm. The design and manufacturing parameters for all the produced scaffolds are resumed in Table 1.
Table 1 – 3DF scaffold geometrical properties and optimised processing parameters.
Scaffold type
Gauge (I.D.) [µm]
Vdep [mm·min-1]
Pres.
[bar]
Temp.
[ºC]
Designed Obtained (*)
Porosity d2 (*)
[µm]
d3
[µm]
d1
[µm]
d2 [µm]
d3 [µm]
1PORP300
G27
(200) 56 5 200 300 150 158±25 285±28 134±26
1.5PORP400 G27
(200) 56 5 200 400 150 152±14 395±38 156±17 46±2%
1.5PORP600 G27
(200) 56 5 200 600 150 187±9 597±14 138±15 66±4%
2PORP600
G25
(250) 196 5 200 600 150 231±28 580±18 148±4
PCW
G25
(250) 196 5 200 400 175 293±27 410±24 152±65 28±10%
G27
(200) 56 5 200 400 150 200±33 402±33 150±12 55±5%
TMp
G25
(250) 56 5 200 - 150 352±32 - - -
(*) Obtained values for d1, d2, d3 and scaffold porosity are presented as mean±standard deviation
9.2.3.2 Electrospinning apparatus
An ES apparatus was used to coat the previously described TMp with a thin PEOT/PBT electrospun meshes (Figure 11c,f). This apparatus is composed of a three-axis system (CNC-STEP, Germany) that allows the positioning and movement of the spinneret during the ES process. The system is also equipped with a syringe pump (KDS 100, KD Scientific) to control the polymeric solution feed rate (F) and a high voltage power supply (Gamma High Voltage Research Inc., USA) capable of generating up to 30 kV. The spinneret (21-gauge blunt tip) was positively charged and the collector was grounded. An aluminium foil was used to cover the collector. The TMp were placed on top of an aluminium foil and the spinneret was aligned with the scaffold prior to the generation of the ES mesh. The environment conditions (temperature of 25 ± 1ºC and humidity of 30%) inside the ES chamber were controlled by means of an ad-hoc ventilation system. A solution of PEOT/PBT (300PEOT55PBT45) was prepared with a concentration of 20% w/v in a solvent mixture of CHCl3 and HFIP (90/10% v/v) as previously reported by Moroni et al [Moroni et al., 2008c]. The production of a thin fibre mesh on top of the TMp was achieved by ES the previously described solution. The collection duration of the electrospun mesh was ranged in 30-120 seconds in order to obtain different mesh thicknesses. A polymeric solution F value of 5 ml·h-1, an applied
voltage (Vapp) of 15 kV, and the distance between spinneret and the collector of 15 cm were kept constant in all the experiments.
Electrospun meshes without TMp were also produced, using a rotating drum as a collector, in order to produce control meshes for biological tests. A rotating drum with a diameter of 60 mm and a length of 60 mm was used to produce thicker meshes and with a constant thickness throughout all the mesh. A rotation velocity of 150 r.p.m. was used in order to collect random electrospun fibres and kept constant during the collection period. Parameters such as electrical voltage, distance between the spinneret and the collector (rotating drum) and the F used for the production of the meshes were the same as previously optimised for the collection over the TMp. The collection time was 30 minutes in order to obtain a membrane with the minimal thickness to be handled and tested in a bioreactor.
9.2.4 Gas plasma treatment
Plasma treatment is a technique that allows the improvement of cell adhesion by changing surface roughness [Woodfield et al., 2006]. All the produced scaffolds were treated with argon (Ar) plasma. Scaffolds were placed inside the radio-frequency glow-discharge chamber (Harrick Scientific Corp., NY, USA). A pre-vacuum was applied until a 0.01 mbar was reached; a subsequent flush with Ar gas was applied four times. The treatment was performed for 30 minutes at a controlled vacuum ranging in 0.1-0.2 mbar with high settings applied to the radio-frequency coil (740V DC, 40 mA DC, 29.6W).
9.2.5 Morphological analyses
9.2.5.1 Scanning electron microscopy (SEM)
Scaffolds architecture and electrospun meshes morphology were analysed by means of SEM (Phillips XL30 ESEM-FEG). Scaffolds and electrospun meshes were sputter-coated with gold (Cressington, Watford, England) for 60 s prior to SEM analysis. SEM micrographs were acquired from the top and cross-section of the scaffolds at different magnifications. For scaffolds cultured with hMSCs, a previous dehydration step of the constructs was performed using series of ethanol/water solutions at increasing ethanol concentrations (70, 80, 90, 96, 100 and 100%) for 30 min each. Samples were dried using a critical point dryer (Balzers CPD-030) with liquid CO2. Scaffolds were sputter-coated with gold (60 s) and cell morphology and newly formed extracellular matrix (ECM) were analysed.
9.2.5.2 Atomic force microscopy (AFM)
AFM was used to evaluate the changes induced by the Ar plasma on the fibre surface topography.
Imaging was performed in Tapping Mode™ with a Multimode™ AFM (Veeco, Santa Barbara, CA) and a Nanoscope 6™ controller operating on tapping mode at room temperature. Typical values for the set-point amplitude ratio varied from 0.7 to 0.9. Scan rates ranged in 0.8-1.0 Hz.
Silicon tips (TESP, Bruker) with a resonance frequency of approximately 350 kHz and scan rate of 1.0 Hz were used. Roughness analysis was performed using AFM software (NanoScope V6, Digital Instrument) on the flattened image height.
Scans of 1µm2 in surface area were taken at three different points and on triplicates of samples. The mean roughness (Ra) and the root mean squared (Rq) were calculated for each scanned area [DeGarmo et al., 2011].
Ra =∑ |yi| (1)
Rq = ∑ yi (2)
where yi is the vertical distance from the centre line.
9.2.6 Porosity
PORP scaffold porosity was assessed by liquid pycnometer measuring the weight and volume of the scaffolds (n=3), and calculated with the following Equation:
P = 1 −× × 100 (3)
where M and V are the measured mass and volume of the scaffold and ρ is the density of the PEOT/PBT (300PEOT55PBT45) equal to 1.2 g/cm3. For the evaluation of the PCW scaffolds the porosity was calculated according to Equation 3, considering the mass of the samples and the volume of the 3D CAD model.
9.2.7 Biological evaluation of PORP scaffolds
The biological evaluation was performed on the PORP scaffolds produced with a scale size of 1.5.
Moreover, two different fibre spacing (d2), 400 µm (1.5PORP400) and 600 µm (1.5PORP600) were investigated.
9.2.7.1 Cell expansion of hMSCs
HMSCs stored in liquid nitrogen were slowly thawed and 4 ml of cold culture medium added.
Centrifugation of cell suspension was performed during 4 min at room temperature and 300 g.
After supernatant removal, cells were resuspended and counted. HMSCs were expanded in six tissue culture polystyrene (TCPS) flasks with 300 cm2 (T300), with proliferation culture medium (CM) [α-MEM, 10% FBS, 0.2 mM ascorbic acid, 2 mM L-glutamin, 100 U/ml penicillin, 100 µg/ml streptomycin and 1 ng/ml basic fibroblast growth factor (bFGF)], until confluence was reached (one week). Cells were collected from the T300 flasks after adding trypsin solution and incubated for 4 minutes in cell culture incubator. A volume of 5 ml of serum-containing medium was added to neutralize trypsin and cells were transferred to a 10 ml falcon and centrifuged for 6 minutes (room temperature, 300 g). The medium and trypsin were removed, proliferation CM was added and cells resuspended for PORP scaffolds seeding.
9.2.7.2 Cell culture of the PORP scaffolds
Prior to seeding, PORP scaffolds were sterilized in a 70% ethanol for 2 h; the process was repeated twice. After sterilization, scaffolds were washed with PBS and placed in 25 untreated well plate, with 2 ml of basic CM (α-MEM, 10% FBS, 0.2 mM ascorbic acid, 2 mM L-glutamine, 100 U/ml penicillin and 100 µg/ml streptomycin) at 37°C and 5% CO2 enriched atmosphere, for one day prior to seeding in order to improve cell attachment. The culture of hMSCs was performed after removing basic CM. An amount of 80 µl of CM containing 0.5x106 cells was added drop-wise on top of each scaffold and incubated for one hour, then volume was adjusted to 2 ml of proliferation CM. Cell culture experiments, under static conditions were performed for 21 days (7 days with proliferation CM followed by 14 days in osteogenic CM or basic CM). Each medium was changed every two days. After 7 days of culture with proliferation CM, the medium was changed to osteogenic CM (α-MEM, 10% FBS, 0.2 mM ascorbic acid, 2 mM L-glutamin, 100 U/ml penicillin, 100 µg/ml streptomycin, 1 ng/ml bFGF and 10-8 M dexamethasone) in order to induce the hMSCs into the osteogenic lineage. Scaffolds were also cultured with basic CM and used as a negative control of differentiation. At each time-point (1, 7, 14 and 21 days), constructs were collected for DNA assay and ALP assay (n=3) and stored at -80ºC. Moreover, samples collected at 7 and 21 days were assayed also by methylene blue staining, histology and SEM analyses.
9.2.7.3 Methylene blue staining
Scaffolds cultured for 7 and 21 days were sliced and a cross-section of each type of scaffold (1.5PORP400 and 1.5PORP600) was analysed. The cell culture medium was removed and scaffolds were washed twice with phosphate buffered saline (PBS). The cells were fixed using 1 ml of formalin at 10% for 30 min. Scaffolds were washed once with PBS to remove formalin. After washing with milli-Q water, the staining with 1% methylene blue (Sigma) was performed by dropping the solution on top of each scaffold. After a short incubation period (15 seconds), scaffolds were washed with milli-Q water until the excess of dye was removed. The prepared scaffolds sections were observed using a Nikon SMZ800 microscope equipped with a digital camera (Qimage RETIGA 1300) and an incident light source (KL1500 LCD, SCHOTT).
9.2.7.4 DNA assay and ALP assay
Quantitative analyses on the cultured scaffolds were performed with DNA assay and ALP assay.
Scaffolds were collected at each time-point in triplicates (n=3) (1, 7, 14 and 21 days) and stored at -80 °C until the end of the cell culture experiments. In order to assess to the largest amount possible of cells present in the scaffold, a pre-cutting of the seeded scaffolds in small sections was performed. Each scaffold was placed in a microcentrifuge tube prior to storage. Before the quantitative analysis, scaffolds were thaw/freeze (30 min at room temperature followed by 30 min at -80ºC) for five cycles. The lysates were prepared after a collagenase digestion of the newly formed ECM. A solution was prepared with collagenase type I and II with a 50:50 (%v/v) in a 90:10 (%v/v) of PBS and collagenase mixture. An amount of 0.75 ml was added to each scaffold and incubated in the dark under shaking at 300 r.p.m. (eppendorf thermomixer confort) at 37ºC for 16 h. For DNA quantification a CyQUANT® NF Cell Proliferation Assay Kit (Molecular Probes, Invitrogen) was used following the manufacturer instructions. An amount of 50 µL of the cell lysates was aliquoted to a 96 well plate and 50 µL of lysate buffer solution RNAse was added and incubated at room temperature for one hour. The dye (100 µl) was added subsequently and incubated during 30 min prior to reading. Plate was analysed using fluorimetric test on a Victor3 plate reader (Perkin-Elmer, U.S.A.). The excitation wavelength used was 485 nm and the emission detection performed at 530 nm.
ALP was assessed by means of biochemical analysis. The lysate of each cell/scaffold construct was stirred, then 10 µl were transferred into a 96 well plate and 40 µl of CDP-star reagent (Roche) was added to each well. After incubation at room temperature for 30 minutes, plate was analysed at 466 nm using chemiluminescence test on a Victor3 plate reader (Perkin-Elmer, U.S.A.).
9.2.7.5 Histology
One PORP scaffold of each type was collected at 7 and 21 days, rinsed twice with PBS and fixed with 10% formalin for 30 minutes at 37ºC. After removal of formalin, scaffolds were rinsed with PBS and stored in 70% ethanol. For the preparation of scaffolds for histology, a successive dehydration procedure with series of ethanol/water solutions at increasing ethanol concentrations (70, 80, 90, 96, 100 and 100%) was performed. A glycol methacrylate (GMA) solution (Merck) was used and scaffolds were embedded according to manufacturer’s instructions. Samples were placed in a Teflon mould and cured at room temperature overnight. Embedded sections of the constructs were cut with a thickness of 10 µm with a microtome and placed on glass slides.
Scaffold slices were stained with Masson-Goldner-Trichrome staining in order to visualize the connective tissue present on the PORP scaffolds. Moreover, Hemotoxylin and Eosin staining was applied to visualize the cells grown inside the scaffolds. After staining, glass slides were kept under a fume hood and let dry. Glass coverslips were used to seal microscope slides with mounting medium (Tissue-tek, Sukuba Kinetek, USA.) and observed under a Nikon E600 microscope.
9.2.8 Statistical evaluation
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.
9.3 Results and Discussion
9.3.1 Scaffold manufacturing and characterization
9.3.1.1 PORP scaffold morphology
Processing parameters were optimised (Table 1) and PORP scaffolds with three different scale sizes (1:1; 1.5:1 and 2:1 ratio) were produced (Figure 5). A lay-down pattern of 0-90º (fibre deposited with a 90º rotation between each deposited layer) was maintained for all the produced scaffolds. The feasibility of the scaffolds with the smallest dimensions (1PORP300) was limited.
The value for d2 was decreased to a minimum of 300 µm, although, the definition of the internal pore architecture was not completely defined and the external geometry was not completely reproduced according to the 3D model (Figure 6a). These scaffolds presented a closed network of pores in some regions and the fusion of the fibres induced by the concentration of temperature
during the production. The values obtained for d1 and d3 were 158±25 and 134±26 µm, respectively (Table 1).
With increasing of the scaffold size, a better definition of the external geometry and internal network of pores was achieved. Since a satisfactory relation between scaffold size and internal pore architecture was achieved for the scaffolds with a scale of 1.5, such scaffolds were further improved and two different scaffolds with d2 of 400 µm and 600 µm (1.5PORP400, 1.5PORP600) were successfully produced (Figure 5, 6b, 7a,b). The values obtained for d1 and d3, with the optimised processing parameters, were 152±14 µm and 156±17 µm for 1.5PORP400, while 187±9 µm and 138±15 µm for 1.5PORP600 (Table 1).
The largest size scaffold produced (2PORP600) presented a well-defined network of pores and it was also possible to produce scaffolds with high reproducibility and accuracy (Figure 5, 6c). In this case, a nozzle of with an I.D. of 250 µm was used and the obtained values for d1 and d3 were 231±28 µm and 148±4 µm, respectively (Table 1).
Figure 5 - Three-dimensional scaffolds for partial PORP with three different scales sizes (1:1, 1.5:1 and 2:1 ratio): a) side view and b) top view.
SEM observations of the prepared scaffolds showed a significant difference in the fibre morphology as presented in the Figures 6 and 7. At higher velocities the fibre resulted not uniform and delamination between subsequent layers occurred. With the optimal parameters, the fibre diameter (d1) obtained should be equal to the nozzle diameter, although this parameter is strongly
influenced by the extrusion of the polymer and the deposition velocity. For scaffolds 1.5PORP400 and 1.5PORP600 produced with a deposition velocity of 56 mm·min-1, the fibre diameter was uniform, although the production time for each scaffold was elevated. The high residence periods of the copolymer inside the heating syringe induced a partial thermal degradation of the PEOT/PBT copolymer and the processing parameters were slightly adjusted according to the changes of polymer viscosity. This was observed also in previously reported work by Moroni et al [Moroni et al., 2006b]. The porosity of the scaffolds produced for a deposition velocity of 56 mm·min-1 varied from 46±2% (1.5PORP400) to 66±4% (1.5PORP600).
Figure 6 – SEM micrographs of PORP-like scaffolds produced by means of 3DF technique: a) 1PORP300, b) 1.5PORP400, and c) 2PORP600. Fibre diameter (d1), fibre spacing (d2) and layer thickness (d3) are reported in the micrograph c). Scale bar = 500 µm.
Figure 7 - SEM micrographs of the PORP scaffold with 1.5:1 ratio: a) 1.5PORP400, and b) 1.5PORP600.
Scale bar = 1 mm. Insert micrographs with detail of fibre spacing (d2), scale bar = 200 µm.
The influence of gas plasma treatment on the surface topography of the scaffolds was evaluated using AFM. As shown in Figure 8 (3D representative images), significant changes on the surface topography of the treated scaffolds were observed when compared to the untreated ones (Table 2).
Untreated scaffolds presented a roughness average (Ra) of 5.8±3.1nm and root mean square (Rq) of 7±3.2nm (Figure 8a). The treated scaffolds presented a higher Ra equal to 14±6 nm and higher Rq equal to 18.1±7.8 nm (Figure 8b).
Figure 8 - Representative AFM images of the surface topography of the scaffolds: a) before plasma treatment, b) after Ar plasma treatment.
Table 2 - Calculated AFM Results of the mean surface roughness (Ra) and root mean square (Rq) of untreated and treated with Ar plasma for 30 minutes 3DF PolyActive™ PORP scaffolds.
Scanned area 1 x 1 µm2
Ra [nm] Rq [nm]
Untreated scaffolds 5.8±3.1 7±3.2
Treated scaffolds 14±6 18.1±7.8
Values are presented as mean ± standard deviation of three independent measurements.
Studies reported by Riekerink et al [Olde Riekerink et al., 2003] showed that Ar plasma treatment of PEOT/PBT films presented a higher surface wettability when compared to untreated films, and this was correlated to the introduction of polar functional groups by reactions of the surface free radicals and oxygen (post-oxidation). The increasing of protein absorption on the surface due to the increase of the wettability results in an enhanced cell adhesion. CO2 gas plasma treatment was also used to improve cell adhesion on PEOT/PBT tissue engineered scaffolds for bone applications [Claase et al., 2003; Claase et al., 2002].
Another example is plasma treated starch/poly(ε-caprolactone) wet-spun scaffolds seeded with human osteoblast-like cell line which showed a cell viability higher than that detected in untreated scaffolds [Tuzlakoglu et al., 2010]. Due to the already recognised influence of the gas plasma treatment on cell attachment, in this study only treated scaffolds were used for cell culture.
Ear ossicle prosthetization has been performed with a wide variety of materials, such as biological grafts, and synthetic or alloplastic materials [Danti et al., 2010; Hildmann et al., 2006]. In some cases the ossicle prosthetization fails and extrusion, lateralization or resorption may occur.
New PORP scaffolds were successfully produced by means of AM technique. To the best of our knowledge, this is the first attempt to produce biodegradable polymeric PORP scaffolds by means of AM techniques. In literature, only one study proposed the production of ossicular replacement prostheses by means of two-photon polymerization AM technique using Ormocer® (organically modified ceramic) and Irgacure® as photoinitiator [Ovsianikov et al., 2007].
A biological validation of the produced scaffolds followed the same model of the previously conducted studies by Danti et al [Danti et al., 2010; Danti et al., 2009], seeding the produced PEOT/PBT PORP scaffolds with hMSCs.
9.3.1.2 PCW scaffold morphology
The production of the PCW scaffolds was performed with the optimised parameters used for the production of the PORP scaffolds. Due to the complexity of the geometry and the reduced thickness, some irregularities on the prepared scaffolds were observed. The lack of structural support when the fibre deposition was performed outside the limits of the previously built layers, hampered an accurate fibre deposition; however, satisfactory scaffolds were produced. A porosity of 28±10% was achieved with the nozzle I.D. of 250 µm (Figure 9a). The PCW scaffold porosity was calculated according to the Equation 1, considering the volume of the scaffold obtained from the 3D CAD model. The values obtained for d1, d2 and d3 with the optimised processing parameters were 293±27, 410±24 and 156±65 µm, respectively.
Some improvement on scaffold pore interconnectivity and porosity was achieved when a nozzle with a smaller diameter (I.D.=200 µm) was used (Figure 9b). In this case a porosity of 55±5% was achieved. The values obtained for d1, d2 and d3 were 200±33 µm, 402±33 µm and 150±12 µm, respectively. The use of a smaller nozzle size or a change in the strategy of the fibre deposition allow the achievement of improved scaffold porosity. When a smaller nozzle size was used, the production rate decreases and the production time of each PCW scaffold significantly increased.
The influence of the porosity and pore size of the PCW scaffolds is not documented in literature.
Notwithstanding, the prosthesis developed and investigated by Blitterswijk et al [van Blitterswijk et al., 1990] showed that, despite of a reduced porosity of the implanted hydroxyapatite scaffolds with 31% of porosity (5% for microporosity and 26% of macroporosity), the pores were filled by new bone. For scaffolds designed for bone applications, the optimal pore size and porosity are not consensual in literature [Karageorgiou and Kaplan, 2005]. It can be expected that a larger pore size and porosity will allow a fast regeneration and vascularization of the implantation site. The implantation site for PCW scaffolds has low demands in terms of mechanical properties; the
scaffolds can have a high porosity level without compromising the structural integrity of the scaffold upon implantation.
Figure 9 – PCW scaffolds produced by means of 3DF technique with different nozzle internal diameter (I.D.):
a) with nozzle gauge 25 – I.D.= 250 µm, and b) nozzle gauge 27 – I.D.=200 µm.
Figure 10 – SEM micrographs of the PCW scaffolds produced by means of 3DF technique with different nozzle internal diameter (I.D.): a) gauge 27 – I.D.= 200 µm, and b) gauge 25 – I.D.= 250 µm. Scale bar =1 mm for all micrographs.
SEM analysis showed a limited porosity and also a limited interconnection of the pore architecture, especially on the surface of the PCW scaffolds produced with the larger nozzle diameter (Figure 10a,b). Due to the small scaffold thickness, an accumulation of polymer on the limits of the pattern originated the fusion of the layers was observed (Figure 10a). Moreover, no porosity was observed on the external surface limiting the pore interconnectivity (Figure 10b). After optimization with a smaller nozzle diameter, completely open porosity was obtained which could be observed both on scaffold cross-section and external surfaces (Figure 10c and 10d, respectively).
One of the first canal wall reconstructions using autologous bone was performed by Merck in 1978 [Mercke, 1987]. Since then, different methods for the reconstruction of the PCW have been proposed, including [Gantz et al., 2010]: osteoperiosteal flaps, composite cartilage/titanium grafts, ceramic alloplasts, costal cartilage, bone cements, and bone pâté.
The produced scaffolds will be validated in future studies following a TE approach in order to aid a fast and biosafe PCW reconstruction. The possibility of producing these types of structures with high accuracy of external geometry and internal pore architecture by means of AM techniques may allow in the future the production of customized implants.
9.3.1.3 TM scaffold morphology
Two different TM scaffolds were designed and successfully produced with a two-step approach combining 3DF and ES techniques (Figure 11). The first step comprised the production of two types of TMp: a) a smaller pattern was prepared with a diameter of 16 mm for static cell culture experiments (Figure 10a,b), and b) a second pattern with a diameter of 25 mm for dynamic cell culture (bioreactor) experiments (Figure 10d,e). The TMp, produced by means of 3DF, was built with a velocity of 56 mm·min-1 and the obtained d1 was 352±32 µm (Figure 12a).
A second step comprised the coating of the TMp with a thin electrospun mesh. An optimization of electrospun mesh morphology and thickness was investigated prior to a coating production step.
Finally, flat electrospun meshes (without TMp) were prepared for future cell culture experiments.
All the TM scaffolds were treated with Ar Plasma treatment for 30 min.
Figure 11 – TM scaffold for dynamic cell culture experiments: a, b and c) 15 mm TM scaffold, and d),,e) and f) 25 mm TM scaffold. a) and d) 3DF deposition pattern, b) and e) TMp produced by means of 3DF technique and c) and f) TMp coated with electrospun mesh for 30 seconds.
For the optimization of the electrospun mesh, a solution of PEOT/PBT (300PEOT55PBT45) with a concentration of 20% w/v in a solvent mixture of CHCl3 and HFIP (90/10% v/v) was prepared as previously reported by Moroni et al [Moroni et al., 2008c]. The processing parameters were optimised (F = 5 ml·h-1, the spinneret-to-collector distance = 15 cm and Vapp = 15 kV) and maintained for all the conducted experiments.
The thickness of the ES meshes varied according to the collection time. For the optimization of the mesh thickness, three times of collection were evaluated (30, 60 and 120 s). Mesh thickness measurements were performed with a micrometre (Mitutoyo Corp., Japan) in three different meshes and three different regions within each mesh. The results of the electrospun mesh thickness are reported in the Table 3.
Table 3 – Electrospun mesh thickness and fibre diameter obtained for different periods of collection.
Processing parameters for the production of ES meshes are also reported.
Polymer concentration 20%v/v
Solvent(s) CHCl3/HFIP (90/10)
Applied voltage (Vapp) [kV] 15
Spinneret to collector [cm] 15
Feed rate (F) [ml·h-1] 5
Collection time [s] 30 60 90 1800(a)
Fibre diameter [µm] 1.8 ± 0.5 (b)
1.9 ± 0.4 (c)
2.0±0.6 2.0±0.8 1.9±0.9
Membrane thickness [µm] 44±19 95±15 146±35 220±56
(a) Collection time performed over the rotating drum, (b) fibre diameter obtained for TMp of 16 mm and, (c) fibre diameter obtained for TMp of 25 mm
According to the thicknesses obtained, a collection time between 30 (44±19 µm) and 60 s (95±15 µm) was selected as it matches the human tympanic membrane thickness range. For this reason, a collection time of 30 s was selected, due to the impossibility to start and stop spinning with accuracy because of the time elapse needed by the spinneret to be translated with the computer controlled system. SEM analyses showed homogeneous fibre morphology for all the produced meshes (Figure 12b-d).
Figure 12 – SEM micrographs of TM scaffold: a) TMp produced by means of 3DF technique, b) TMp with 16 mm diameter coated during 30 s with electrospun fibres (Inset image on b) refers to the electrospun fibre detail, scale bar of 10 µm, c) TMp with 25 mm diameter coated during 30 s with electrospun fibres, d) obtained electrospun mesh collected on a rotating drum during 30 min.
The collected ES mesh above the 16 mm TMp was composed of ultrafine fibres with a diameter of 1.8 ± 0.5 µm (Figure 12b) and 1.9 ± 0.4 µm above TMp with 25 mm (Fig 12c). This small variation on the fibre diameter was not significant (p > 0.05).
One of the first studies conducted for the production of TM scaffolds with PEOT/PBT copolymer was performed by Bakker et al [Bakker et al., 1990a; Bakker et al., 1990b]. PEOT/PBT scaffolds were produced with a thickness of 100 µm, porosity of 50% and pore size of 160 µm [Grote et al., 1991]. The PEOT/PBT biocompatibility was tested on a rat model and observations performed
after two and four weeks showed epidermis and epithelium layers covering the implant, with mild foreign body reaction after implantation. After one year more than 50% of the copolymer was degraded.
Ideally, the production of the TM scaffolds could be achieved only by the production of different layers of ES fibres resembling the natural ECM of the human TM. Although the limited thickness of the ES mesh needed (human TM thickness ~ 100 µm) could limit their application and handling upon biological in vitro tests (dynamic and static) and finally upon surgery. To overcome this limitation, the combination of 3DF and ES was performed. The presence of the TMp can aid in diverse ways: a) avoid of the curling of the graft due to the small thickness of the electrospun mesh, b) the external contour of the TMp can aid fixation to the TM bony annulus, c) the centre part of the TMp can aid fixation to the remain malleus or to an implanted PORP scaffold, and d) avoid the perforation of the TM scaffold.
Generally, the combination of 3DF and ES techniques is proposed in literature because of the possibility of producing scaffolds with good mechanical properties, due to the macro and micro size 3D structure produced by AM techniques and to the good biological performance for the presence of nano-sized electrospun fibres [Park et al., 2008] (see chapter 3). Moreover, it has been shown that the presence of oriented ultrafine fibres can induce different cell morphology [Nain et al., 2008] (see Chapter 3). In the case of the produced TM scaffolds we expect that the presence of the TMp can aid to the reorganization of the cells into an arranged pattern and ideally to produce matrix following the natural human TM ECM fibre arrangements. Moreover, upon implantation using the acellular approach, the presence of the ES mesh can favour the migration of cells from the surrounding tissue.
Electrospun meshes without TMp were also produced using a rotating drum as a collector in order to achieve a uniform mesh thickness. The electrospun meshes were collected during 30 min in order to obtain a minimum thickness suitable to be handled. The electrospun meshes produced were composed of ultrafine fibres with a diameter of 1.9±0.9 µm, a thickness of 220±56 µm and a porosity of 80±0.8%. These meshes were produced as controls in order to compare the hybrid (3DF+ES) TM scaffolds from a biological point of view; these studies are part of ongoing work.
Several are the well-established techniques applied worldwide for TM grafting. The most common techniques rely on the autologous graft transplantation (e.g. postauricular fascia or cartilage tissue).
Regarding the application of the fascia tissue, so called outer surface techniques, it comprises several steps with some complexity, although good results are obtained [Fayad and Sheehy, 2010].
Another approach uses cartilage tissue, and two techniques are generally employed: the perichondrium/cartilage island flap, which uses tragal cartilage, and the palisade technique, which uses cartilage either from the tragus or cymba [John L, 2010]. Some concerns regarding the conductive hearing loss due to the cartilage thickness and rigidity are found in literature [Dornhoffer, 1997; Kerr, 1980], although some reports demonstrate also effective results [John L, 2010; Yung, 2008].
Satisfactory results are usually obtained when autologous tissue is used, even if the long and complex surgical procedures may limit in some ways the positive results after surgery. Despite of the low site morbidity, the risks of additional incisions and the lack of tissue in the vicinity of the perforated area for recurrent patients may limit the use of autologous tissues [Kaftan, 2010]
Homografts were also proposed [Kerr, 1980; Marquet, 1971], but the risk of disease transmission limits the use of this approach. Decellularized xenografts have been also studied, although only some in vivo studies on animal models were performed [Deng et al., 2009].
Alloplastic materials have been also proposed for TM regeneration. Since the complex harvesting process and preparation of grafts is avoided, the use of allopastic materials leads to a reduction of surgery procedure duration.
Strategies applying scaffolds intend to achieve a faster treatment of the TM perforations and generally two approaches are studied: the first approach uses an acellular scaffold with or without the necessity of surgery [Kanemaru et al., 2011; Kim et al., 2011], while a second approach involves an in vitro pre-culture of cells before implantation [Ghassemifar et al., 2010; Levin et al., 2010]. Recently, the combination of bFGF and gelatine scaffolds was proposed to regenerate perforated TM and healing was observed in more than 98.1% [Kanemaru et al., 2011]. Despite of the remarkable rates of closure of the TM, in some cases up to four treatments were necessary and in one patient the TM closure failed. The failure cause was not reported, and to ensure the success of this regenerative treatment, only patients with dry TM perforations and with absence of infection during 3 years were considered [Kanemaru et al., 2011]. Another study suggested the topical application of platelet-derive growth factor (PDGF) [Roosli et al., 2011]. In this case no scaffold was used and results obtained after the application of PDGF were not satisfactory.
With the new developed TM scaffolds we expect to have a faster regeneration of the TM perforations and the presence of multiscale features may induce cells to a faster ECM generation, capable of mimicking the physiological function of the native TM. Despite of the large amount of techniques developed in the last decades, an optimal procedure for TM regeneration is not consensual yet. Novel tissue engineered TM scaffold can aid and improve results of surgery procedures. Moreover, the possibility of providing surgeons with off-the-shelf products ready to be implanted can simplify surgical procedures such as the previously mentioned ones involving autologous tissue. The developed TM scaffolds can enhance the success rate of TM regeneration either applied pristine (i.e., acellular) or after an in vitro cell culture step (TE approach).
The possibility to produce a patient-specific TM is also possible due to the low production time necessary to manufacture these types of structures. The biological evaluation of the manufactured TM scaffolds is part of the ongoing work.
9.3.2 Biological results of PORP scaffolds
In order to validate suitability of the scaffold internal architecture and pore size for bone regeneration, scaffolds were seeded with hMSCs at a density of 0.5x106. Cell culture studies were conducted for 21 days: the first 7 days with proliferation CM and the following 14 days with osteogenic or basic CM. Scaffolds were collected at different culture time-points and processed for methylene blue staining, histology, SEM and DNA and ALP assays.
9.3.2.1 Evaluation of cell distribution by means of methylene blue staining
Methylene blue staining was used to evaluate the distribution of the hMSCs cultured on the scaffolds at 7 and 21 days (Figure 13). Images showed that a large amount of cells adhered on the exterior surfaces of both type of scaffolds when compared to the inner regions. No significant differences were observed on the two time points of observation. It was also possible to observe in the cross-section of the scaffolds that cells were mainly localised on the fibre-fibre intersection regions.
Figure 13 - Optical microscope images of PORP cross-sections and external surfaces stained with 1%
methylene blue after cultured with hMSCs: a) 7 days cell culture performed with proliferation CM, b) 21 days cell culture performed with osteogenic CM (7 days with proliferation CM followed by 14 days with osteogenic CM). Scale bar = 500µm.
9.3.2.2 DNA assay
Cellularity or cell number was consistent throughout the cell culture experiments (Figure 14).
Figure 14 – Quantitative analysis of the cellularity observed on scaffolds seeded with hMSCs at different time-points. All the samples were collected in triplicate (n=3) and analysed.
Cellularity was determined according to the estimated value (7.18 pg/cell) for the nuclear DNA content of a human diploid cell [Taylor et al., 1989]. The average DNA amount detected after 24 hours of cell seeding was 42% for 1.5PORP400 and 40% for 1.5PORP600 scaffolds (Figure 14a).
Results showed a minor difference between scaffold types, although it was not statistically significant. The high cell loss upon seeding is generally related to the low scaffold retention capabilities. Moreover, the initial cell adhesion is also correlated to different factors like pore size, porosity, pore architecture, surface topography and chemistry [Danti et al., 2010; Danti et al., 2009;
Solchaga et al., 2006]. The big standard deviation observed in the cell amount between scaffolds of the same type, for each time-point, might be correlated to the high volume used for cell seeding (80 µl). Further optimization should be performed in order to improve seeding uniformity between scaffolds of the same type. This can be achieved by tweaking the volume seeded evaluating the optimal amount that a scaffold can retain [Danti et al., 2010]. The cell number observed after 7 days of culture in proliferation CM was slightly higher than that observed after day one, although
the changes were not statistically significant. The amount of cells present after 7 days in proliferation CM were ~ 0.254x106 for 1.5PORP400 scaffold and ~ 0.220x106 for 1.5PORP600 scaffold.
9.3.2.3 ALP assay
HMSCs can be induced to differentiate towards different cellular phenotype (e.g. osteoblasts, chondrocytes and adipocytes), when cultured in specific differentiation media [Li et al., 2005]. In this study, an osteogenic medium containing dexamethasone was used to induce hMSCs towards the osteogenic lineage. ALP is an early marker of osteogenesis [Danti et al., 2010] and it was used to evaluate the differentiation activity of the hMSCs cultured in the PORP scaffolds. After 7 days of cell culture with proliferation CM, the PORP constructs were subsequently cultured with basic or osteogenic CM. Results showed a higher ALP activity when cultured in osteogenic CM when compared with basic CM, determining a higher osteogenic differentiation in osteogenic CM (Figure 15).
Figure 15 – Alkaline phosphatase (ALP) activity analyses of cell cultured scaffolds with: a) basic CM (7days with proliferation CM followed by 7 and 14 days with basic CM), b) osteogenic CM (7days with proliferation CM followed by 7 and 14 days with osteogenic CM).
The ALP values measured for the scaffolds cultures with osteogenic CM were significantly higher in both time-points when compared with the scaffolds cultures in basic CM. Moreover, a significant up-regulation of the ALP activity was observed on the scaffolds in osteogenic media at 21 days when compared to 14 days in both type of scaffolds (1.5PORP400 and 1.5PORP600). No significant difference was observed when comparing the two types of scaffolds for the same time- points and the same CM, revealing that the difference in porosity and pore size seems not to affect the level of ALP production and the osteogenic differentiation in basic and osteogenic CM.
9.3.2.4 Histology
Hematoxylin & Eosin staining was used to observe cell distribution on the seeded scaffolds. After 21 days (Figure 16), a superior amount of cells were observed on the seeded scaffolds when compared to the scaffolds cultured for 7 days. Some clusters of cells were also observed on the bottom of the 1.5PORP400 scaffolds (Figure 9c). A significant amount of cells was also observed on the exterior surfaces of the constructs in accordance with what was observed using methylene blue assay.
Figure 16 - Histological images of the PORP scaffolds seeded with hMSCs stained with eosin and hematoxylin at cell culture endpoint (7 days of proliferation CM followed by 14 days of osteogenic CM): a) 1.5PORP400 (white arrow points out the cell clump) and b) 1.5PORP600. Scale bar = 100 µm.
Masson-Goldner-Trichrome staining was used to observe the presence of collagen fibres. In both the constructs, collagen fibres, stained in blue/green, appeared well organized and arranged around to the scaffold material. However, a superior amount of fibres was observed on the 1.5PORP400
scaffolds cultured for 21days when compared to 1.5PORP600 scaffolds (Figure 17). Cell clusters were also identified as previously described.
Figure 17 - Histological images of the PORP scaffolds seeded with hMSCs stained with Masson- Goldner- Trichrome at cell culture endpoint (7 days of proliferation CM followed by 14 days of osteogenic CM): a) 1.5PORP400 (white arrow points out the cell clump) and b) 1.5PORP600. Scale bar = 100 µm.
9.3.2.5 SEM analysis of cell morphology
SEM micrographs of the cell cultured scaffolds showed a significant amount of cells covering the fibres after 7 days. A lower amount of cells was detected in the interior of the scaffolds (cross- sections micrographs) in accordance to what previously described for methylene blue staining (Figure 13). The limited migration of cells towards the scaffold interior is largely reported in the literature and is correlated to a limited diffusion of nutrients in static culture experiments [Salerno et al., 2009; Shi et al., 2007]. A larger amount of cells and newly formed ECM were observed on the external surfaces of both types of scaffolds when compared with the internal regions (Figure 18). It was also possible to observe a cell layer covering the scaffold fibres after 7 days of culture.
After 21 days of culture, an increase in cell number and ECM was observed both in the interior and exterior regions of the scaffolds when compared to 7 days. Moreover, a cell sheet covered the scaffold external surface closing completely the pore region (Figure 18).
Figure 18 – SEM micrographs of the 1.5 PORP400 and 1.5PORP600 scaffolds seeded with hMSCs during 7 (proliferation CM) and 21 days (7 days of proliferation CM followed by 14 of osteogenic CM): left column with cross-sections (scale bar = 1 mm) and right column with the external surfaces (scale bar = 200 µm).
Inserts in the left side column represent detail regions of the cross-sections (scale bar = 100 µm)
