Contact us

My IOPscience

Coaxial additive manufacture of biomaterial composite scaffolds for tissue engineering

This content has been downloaded from IOPscience. Please scroll down to see the full text. 2014 Biofabrication 6 025002 (http://iopscience.iop.org/1758-5090/6/2/025002) View the table of contents for this issue, or go to the journal homepage for more

Download details: IP Address: This content was downloaded on 17/02/2015 at 14:16

Please note that terms and conditions apply.

Biofabrication Biofabrication 6 (2014) 025002 (9pp)


Coaxial additive manufacture of biomaterial composite scaffolds for tissue engineering R Cornock, S Beirne, B Thompson and G G Wallace Intelligent Polymer Research Institute, ARC Centre of Excellence for Electromaterials Science, University of Wollongong, NSW 2522, Australia E-mail: [email protected] Received 16 May 2013 Accepted for publication 9 January 2014 Published 21 March 2014 Abstract

An inherent difficulty associated with the application of suitable bioscaffolds for tissue engineering is the incorporation of adequate mechanical characteristics into the materials which recapitulate that of the native tissue, whilst maintaining cell proliferation and nutrient transfer qualities. Biomaterial composites fabricated using rapid prototyping techniques can potentially improve the functionality and patient-specific processing of tissue engineering scaffolds. In this work, a technique for the coaxial melt extrusion printing of core-shell scaffold structures was designed, implemented and assessed with respect to the repeatability, cell efficacy and scaffold porosity obtainable. Encapsulated alginate hydrogel/thermoplastic polycaprolactone (Alg-PCL) cofibre scaffolds were fabricated. Selective laser melting was used to produce a high resolution stainless steel 316 L coaxial extrusion nozzle, exhibiting diameters of 300 μm/900 μm for the inner and outer nozzles respectively. We present coaxial melt extrusion printed scaffolds of Alg-PCL cofibres with ∼0.4 volume fraction alginate, with total fibre diameter as low as 600 μm and core material offset as low as 10% of the total diameter. Furthermore the tuneability of scaffold porosity, pore size and interconnectivity, as well as the preliminary inclusion, compatibility and survival of an L-929 mouse fibroblast cell-line within the scaffolds were explored. This preliminary cell work highlighted the need for optimal material selection and further design reiteration in future research. Keywords: additive fabrication, additive biofabrication, co-axial extrusion, tissue engineering, cell encapsulation, alginate, polycaprolactone (Some figures may appear in colour only in the online journal)

the wound contraction is blocked and instead regeneration is induced [4]. There are considerable challenges associated with this; the developed tissue replacement must be durable [5], nonimmunogenic (or autologous) in nature [6, 7], and exhibit the same function and morphology as the target tissue [3]. Successful bioscaffolds must also demonstrate desirable structural strength, dimensions, porosity and pore size [8, 9], as well as pore interconnectivity and suitable bioactivity which reflect that of the body’s natural ECM to guide cell development and proliferation [10]. Furthermore, the scaffolds fabricated are generally required to exhibit specific structural strength, dimensions and porosity, as well as pore interconnectivity and bio-affinity. Bioscaffolds were reviewed

1. Introduction Tissue engineering (TE) is an emerging biomedical discipline which aims to create artificially engineered products of natural or synthetic biomaterials, often coupled with the strategic placement of cells [1], into three-dimensional networks or bioscaffolds. These bioscaffolds should mimic the physiological and geometric architecture of the body’s extracellular matrix (ECM), and hence have a potential to replace damaged or defective tissue, or even entire organs [2, 3]. In the event of major injury, the physiological response of the body is one of inflammation, followed by cell mediated wound contraction and the generation of scar tissue in repair [4]; however if a suitable ECM analogue can be implemented, 1758-5082/14/025002+09$33.00


© 2014 IOP Publishing Ltd

Printed in the UK

R Cornock et al

Biofabrication 6 (2014) 025002

[30]. However this technique requires the use of barriers within the sample to segregate areas of different cell function, which would restrict the total interconnectivity of the produced scaffold. Here we have developed a coaxial melt extrusion printing (CMEP) system capable of encapsulating low stiffness alginate hydrogels with higher modulus polycaprolactone, a thermoplastic biopolymer, into core-shell configurations. The produced architectures were assessed with respect to their reproducibility, the locale of the core material with respect to the shell cladding (offset), scaffold porosity, the encapsulated volume fraction achievable and initial cell viability. This method could later be integrated into multi-head deposition systems.

in Langer and Tirrel [11], Muschler et al [12] and Hollister [13]. Many hydrogels such as those derived from alginic acid (sodium alginate) demonstrate the ability to integrate rapidly into the ECM of the body and facilitate the passive diffusion of oxygen and mass transport, as well as exhibiting desirable degradation qualities [2]. These macroporous hydrogel solutions also boast the ability to encapsulate cells, growth factors and proteins and are easily processable using fibre spinning and solid freeform fabrication techniques [14]. As such, alginate has been utilized in wound dressings, cell encapsulation and primitive TE applications previously [15]. While hydrogels present ideal properties for soft tissue applications, many hard tissue applications such as bone and musculoskeletal analogues require greater strength to maintain their structure and resist the shear or torsional forces they are subject to following implantation [16]. Subsequently, composite scaffolds have been implemented which consist of an organized thermoplastic biopolymer scaffold which provides strength combined with cell-seeded hydrogels [14]. One such thermoplastic biopolymer is poly-εcaprolactone (PCL). PCL demonstrates good melt processability [17, 18], with a low melting temperature ranging between 56–65 C [17, 19–21], and a high decomposition temperature of 350 ◦ C [22]. It also has sufficient tensile strength (σ T, 4–785 MPa), Young’s modulus (0.19– 0.44 GPa) [18, 23, 24] and percentage elongation at break (20–1000%) [19, 24] to make it an ideal material for bioscaffold support fabrication. The preparation of PCL as porous, biocompatible, biodegradable structural scaffolds (often loaded with therapeutics [25]) through extrusion printing and fibre spinning techniques has been identified and researched extensively in recent years [26–29]. While polycaprolactone has been observed to support cell adhesion, exhibit both hard and soft tissue compatibility and tuneable porosity, its low number of cell recognition sites, hydrophobicity and extended degradation period limit the applicability of the polymer as a stand-alone bioscaffold. Consequently, there has been much investigation into the use of polymer blends and hybrid or composite bioscaffolds to meet the expansive requirements of specific TE scenarios [21, 30]. There exist many processing techniques for the fabrication of functional TE scaffolds, such as solvent casting, gas foaming, emulsion freeze drying and fibre spinning [31, 32]. However each of these methods has inherent limitations in either complexity of structure, scaffold thickness, pore interconnectivity or shape [6]. Consequently, solid freeform fabrication (SFF), a form of additive manufacture (AM), has become a dominant method for the quick design and reproducible fabrication of TE architectures as it provides an ordered and customizable deposition, can integrate cells directly into the sample, and supports the use of multiple materials using multi-head hybrid bioprinters. For such bioplotting systems, thermoplastic polymer is first deposited and the resulting ‘channels’ filled with hydrogel [30]. The resulting constructs tend to assume the mechanical properties of the thermoplastic polymer, and cell viability after three days has recently been demonstrated using such techniques

2. Method and materials 2.1. Material preparation

Alginic acid with sodium salt isolated from brown algae (lot #090M0092V) was obtained from Sigma-Aldrich. 2% wt/wt alginate solutions were produced through the addition of alginic acid to Milli-Q water under constant stirring and for 30 min. For the pre-crosslinked samples used in the cell study, CaCO3 (BDH Chemicals, lot #38923) was added to the Milli-Q water to a concentration of 60 mM and 15 μL mL−1 gluconolactone added, causing a weak hydrogel to form after 45 min. For increased visibility in coaxial structures, methyl red was added in low concentrations to selected solutions for imaging. PCL of molecular weight ∼80 kDa was obtained from Sigma-Aldrich (lot #07327ME) in pellet form and used as received. 2.2. Coaxial melt extrusion printer design

A dual material coaxial nozzle approach was designed and fitted as described in figure 1. Two PTFE-shielded heat jackets enclose the material reservoirs. These reservoir barrels are oriented at 60 ◦ C and consist of a stainless steel chamber with a standard male luer lock fitting at one end and a pressure coupling at the other. The pressure couplings provide an inlet from a variable pressure gauge supplied with nitrogen gas. The extrusion nozzle itself contains two separated chambers for the individual inputs, and at the intersection in the centre of the component the interior chamber sweeps to within the exterior, aligning ultimately at the extrusion tip (figure 2). The interior nozzle protrudes 200 μm from the exterior nozzle. This was designed using Dassault Systemes Solidworks 2012, and the FloExpress extension was used to model flow behaviour and optimize the design. Selective laser melting (SLM) was utilized in the fabrication of the coaxial melt extrusion nozzle, using Technile Spezialpulver refined stainless steel 316 L powder (particle size 23–60 μm) as both the build and supporting material. The supporting material was later removed before electropolishing. A Struers Lectropol-5 Electrolytic Polishing Control System was used to electropolish the extrusion nozzle for a period of 120 s. An apparatus as described in figure 3 was designed and utilized for the electropolishing procedure. 2

R Cornock et al

Biofabrication 6 (2014) 025002

Figure 1. Schematic of the CMEP system, including dual material reservoirs, with independent temperature and pressure controls.



(C )

(D )

Figure 2. Coaxial melt extrusion tip design. (A) Isometric and (B) section drawings. (C) Example of flow analysis using FloExpress extension for Solidworks 2012. (D) 3D printed prototype produced using an Objet Connex 350 in MED610 proprietary transparent material. 2% Alginate hydrogel solution with methyl red (left) and brilliant blue (right) are loaded into the separate chambers to show the internal architecture of the component.

for investigation with varied process parameters as described in table 1. Single coaxial fibres were also printed at a length of 60 mm for characterization.

The electrolyte bath used consisted of 45 mL phosphoric acid (85 wt%), 30 mL sulfuric acid (95 wt%) and 25 mL distilled water per 100 mL total volume. A 25 mm length of platinum wire was utilized as the cathode, while the sample itself acted as the anode. To obtain a uniform polish, the nozzles were slowly rotated during electropolishing.

2.4. Fibre and scaffold characterization

Fibre volume and volume fraction were observed through length and average radius measurements obtained using a LEICA M205A optical microscope. Measurements of the fibre diameter were made by orienting the sample vertically. Measurements of the fibre length and diameters were made in triplicate. With the use of indicators within the inner material, a good representation of coaxial fibre morphology could be elucidated using optical microscopy. The volume fraction could then be used in conjunction with the material density to determine the volumetric flow rate of the extrusion. LEICA microscopy was also used to evaluate the percentage offset of the core fibre from the centre of the coaxial fibre. Scanning electron microscopy was carried out using a JEOL JSM-7500FA Field Emission Scanning Electron Microscope. Samples were first sputter-coated with Au particles using an Edwards sputter coater. The mechanical characteristics of cofibres were observed using a Shimadzu EZ-L Universal Mechanical Tester. Fibres were cut to a length of 20 mm, a gauge length or clamp spacing

2.3. Coaxial melt extrusion printing of scaffolds

An adapted Korean Institute for Machinery and Materials Scaffold Plotting System (SPS1000) was used to generate coaxial scaffolds. The feed rate, extrusion pressure and temperature process parameters were manually optimized to adjust for the individual material viscosity and temperatures required to induce flow as previously observed in the rheological study [33]. As both materials were found to be pseudoplastic or shear-thinning (p < 1.00), the first print line was removed from the study to eliminate error caused by the initial shear-thinning of the materials. Korean Institute of Machinery and Material (KIMM) Bioplotter software was used to slice primitive lattices and produce a G-code tooling path for fabrication. For CMEP it was also found to be beneficial to the final outcome to run the outer extruder 1–2 s prior to the interior extruder to ensure the core material was completely encapsulated. Alginate/PCL coaxial scaffolds were fabricated 3

R Cornock et al

Biofabrication 6 (2014) 025002


(B )

(C )

(D )

Figure 3. (A) Setup for the electropolishing of printed components. Includes a platinum counter-electrode, sample holder, magnetic

stirrer/hotplate for stirring of electrolyte solution during polishing to disperse dissolved ions and an electrolyte solution containing distilled water, phosphoric acid and sulfuric acid. (B) SLM fabricated CMEP nozzles with support material. (C) KIMM SPS 1000 fitted with the fabricated CMEP nozzles and in operation. (D) LEICA optical microscopy image of the coaxial extrusion nozzle tip following 120 s electropolishing at 5.0 V. Table 1. Range of temperatures, pressures, nozzle diameters and feed rates observed during preliminary optimization of the CMEP process.

Reservoir material Alginate Polycaprolactone

Nozzle diameter 200–300 μm 900–1100 μm

Feed rate −1

100–300 mm min 100–300 mm min−1

of 10 mm was used and maximum loading of 20 N applied with a strain rate of 1 mm min−1 . Strain was observed as a function of stress for an applied and increasing force, and the tensile strength, Young’s modulus and elongation at break were determined. Samples were measured in triplicate. The porosity of printed coaxial scaffolds was observed through gravimetric determination, in accordance with equations (1) and (2). The observed volume fraction for the alginate-PCL cofibres was utilized in establishing the bulk material density. The pore size was then determined for each construct using optical microscopy ms ds = (1) As Hs   ds . (2) Ps = 1 − dB Where ms , As and Hs are the mass, area and height of the scaffold respectively, ds and dB are the scaffold and bulk material densities respectively, and Ps is the scaffold porosity.

Applied pressure


1–25 kPa 400–600 kPa

23 ◦ C 90–140 ◦ C

2.5. Cell study

A mouse fibroblast cell line (L-929 cells) was grown to 80% confluence in a T-75 flask.The cells were trypsinized and centrifuged at 500 g for 5 min to pellet the cells and resuspended in 300 μL Dulbeccos Modified Eagles Medium supplemented with 10% foetal bovine serum to a concentration of 2 × 107 cells mL−1 . 2 μL of 1 mg mL−1 propidium iodide (Sigma-Aldrich) and 2 μL of 1 mM calcein AM in DMSO (Invitrogen) were added to the cells, and incubated at 37 ◦ C for 20 min. The cells were then added to the gels, with 300 μL diluted into a prepared alginate solution, before addition of the gluconolactone and calcium carbonate to a final volume of 15 mL for an approximate final cell concentration of 4 × 105 cells mL−1 . The live–dead stained cell-laden gels were observed using confocal microscopy in three configurations: as cast, after manual extrusion into printed hollow fibres and after CMEP with PCL as a shell material. 4

R Cornock et al

Biofabrication 6 (2014) 025002




Figure 4. (A) Optical microscopy image of CMEP dual nozzle, designed at 300 μm/900 μm. (B) SEM image of an SLM produced surface,

and (C) SLM surface following 5 min of electropolishing at 5.0 kV with 150× magnification and a working distance of 8.0 mm.

(A )

(B )

Figure 5. Process parameter optimization for geometry minimization. (A) Alg-PCL coaxial fibre diameter as a function of applied pressure to exterior PCL reservoir. (B) Alg-PCL coaxial fibre diameter as a function of temperature of PCL reservoir.

as 0.60 mm and an alginate volume fraction as high as 0.45. The volume fraction could be modulated through variation in the pressure applied to the alginate material reservoir, as the exterior PCL extrudate was observed to swell with alginate volumetric flow rate to the point of rupture. The grid size of the scaffolds was selected such that a theoretical 200 μm spacing would exist between adjacent fibres based on fibre width observation. In practice, the fibre thickness and consistency of successive layers changes dramatically from that of the first layer. Some examples of the coaxial scaffolds produced and the coaxial extrusion process are shown in figure 6, along with the morphology of the surface, cross-section and interior walls of these scaffolds. While most of the scaffolds show a relatively consistent surface, there were instances of voids developing in the samples, likely due to residual solvent from the cleaning of components affecting the consistency of the polymer melts. The scaffolds show a definitively flattened side, caused by the initial adhesion to the glass substrate.

3. Results and discussion 3.1. High resolution coaxial print components using selective laser melting

SLM was successfully used to produce a functional coaxial extrusion nozzle component. The nozzle was designed to harbour low diameter nozzles (300 μm/900 μm for the interior/exterior respectively) and electro-polished to improve the surface finish and hence flow homogeneity (figure 4). Following initial experimentation, the design was reiterated to encourage an even flow distribution around the outer nozzle. 3.2. Produced structures and morphology

Following initial successes in coaxial fibre extrusion, printing pressures (figure 5(A)), temperatures (figure 5(B)) and feed rates were optimized using the observed flow rate and fibre diameters produced, as well as previously reported material rheology and internal thermal gradient information [33]. This was vital in improving the level of geometry minimization possible in the coaxial constructs. The CMEP system was successfully used to produce Alg/PCL composite fibres and 2–3 layer square lattice scaffolds using the parameters described in table 1. Staining of the alginate hydrogel using methyl red prior to extrusion demonstrated that minimal mixing occurred in the samples between the PCL and the gel. The individual fibres produced exhibited diameters of as low

3.3. Printed hollow fibres allow post-process core-loading

Further investigations revealed that, using Milli-Q water and an appropriate gauge of syringe, the alginate core material could be manually flushed from the PCL, leaving hollow PCL fibres of similar dimension. As such, we were able to replace the core material of the fibres post-process with, for example, cell-loaded hydrogel with no heat exposure. These printable 5

R Cornock et al

Biofabrication 6 (2014) 025002

(A )

(B )

(C )


(E )

(F )

(G )

(H )

Figure 6. (A) CMEP of Alg/PCL cofibres. (B) Tensile testing of printed coaxial fibres. (C)–(E) Multilayered Alg-PCL scaffolds, showing

extensive pore interconnectivity. This interconnectivity was qualitatively measured using optical microscopy through the XY, XZ and YZ planes of each scaffold. (F) Optical microscopy image of a printed hollow fibre produced through post-process removal of alginate from a coaxial fibre. (G), (H) SEM images of a hollow fibre, showing the surface morphology of the interior (scale bars each represent 100 μm).

channels may be of interest in further applications such as in biocompatible fluidics systems. While hollow PCL fibres could also be fabricated by supplying air and water through the interior nozzle, the best results were achieved when a place holder such as alginate was used and flushed post process.

bioprinter. The highest porosity the coaxial scaffolds exhibit reaches 68%. This is comparable with the observed values for the single material PCL reference scaffold (76%). Although geometries with porosity of up to 75% tend to increase, for example, the ingrowth of bone tissue into the scaffolds [34], this is also highly dependent on pore size, which can be measured via microscopy means. These measurements revealed a 0.54 ± 0.03 mm pore size of Alg-PCL scaffolds. This is significantly larger than scaffolds in development for bone TE applications (pore size 350 μm [34]). The pore size here is limited by the diameter of the extruded fibres. By reducing the fibre diameter, we could increase the number of pores per unit area and thus observe a smaller pore size for a given porosity. Further design iterations and optimization of the CMEP process would allow for the tailoring of coaxial scaffold porosity and pore size for specific TE applications.

3.4. Coaxial scaffolds are reproducible

An important consideration regarding the reproducibility of coaxial fibres and scaffolds is the core fibre offset. Ideally, the core material will be positioned at the centre of the cross-section of the exterior material. Initially the percentage offset of the alginate within the PCL was quite high (>40%), however after reiteration of design, reduction of the inner nozzle protrusion to 200 μm and increasing the feed rate of the apparatus (220 mm min−1 ), offsets as low as 10% were observed (figure 7). The highest resolution scaffold generated exhibited a layer height of 600 μm, a grid size (distance between fibre centres) of 950 microns, and an alginate volume fraction of 0.37. The fibre average fibre offset for this scaffold was found to be 15.82 ± 2.8% (n = 5).

3.6. Mechanical characteristics

The mechanical properties of printed Alg-PCl coaxial fibres were measured using tensile testing to record the tensile strength (σ ), elastic modulus (E) and elongation at break (break ) of the composite (figure 8). These properties were evaluated for a consistent alginate volume fraction of ∼0.4 and compared with literature values for the pure materials. The tensile tests yielded a tensile strength of 2.8 ± 0.3 MPa,

3.5. Coaxial scaffold porosity is tuneable

The scaffold porosity was determined for the highest resolution Alg-PCL coaxial scaffolds produced, and compared to typical non-coaxial scaffolds produced using the unmodified 6

R Cornock et al

Biofabrication 6 (2014) 025002

(A )

(B )

Figure 7. (A) Optical microscopy image and demonstration of cross-sectional measurements for an Alg-PCL cofibre. Interior and exterior

cross-sectional areas were divided into equal sections and their centre distance (green) measured to determine the offset percentage. (B) Core material percentage offset as a function of feed rate between 200–300 mm min−1 for Alg-PCL cofibres. Overlayed in the graph are example optical microscopy images at each data point.

on survival of the L929 fibroblast cell line. The 2% alginate solutions with CaCO3 , GL and cell additions were divided into three groups. While one gel was cast onto a microscope slide as a control to ensure the cells were compatible with the gel mixture, the second was injected into hollow printed PCL fibres, to observe cell response to shear stresses alone, and a third extruded into a PCL fibre using CMEP to observe the effects of temperature on cells during extrusion. Figure 9 shows the results of these studies utilizing confocal microscopy to image the live/dead stained samples. The results indicate that the cells remained alive for at least 2 h in the 2% alginate hydrogels (viability at 2 h of 93.0 ± 0.5%), and that the shear stress applied when extruding through a 24 gauge needle (selected to mimic the inner diameter of the extrusion tip) did not significantly damage the cells (viability at 2 h of 89.2 ± 0.3%). However the coaxial extrusion of cell-laden alginate solutions inside the PCL melt shell material was found to stress cells to the point wherein virtually no metabolic activity could be observed (viability after 2 h of 0.7 ± 0.7%), with all cells showing compromised membrane integrity by live/dead staining. This suggests that heat stress was the cause of cell death. The coaxial fibres (or bundles thereof) as well as scaffolds produced lend themselves to linear TE applications such as in muscle and nerve repair, (providing a macroporous coaxial biomaterial scaffold without the need for cell inclusion), as well as in enabling the multi-phase release of therapeutics. However, several steps should be taken in future work if the viability of cell inclusion during printing is required. The system should be adapted to include a third temperature control at the extrusion nozzle, thus removing the thermal gradient which occurs between the melt reservoirs and the nozzle, reducing cell exposure to elevated temperatures. The input temperature could then be lowered significantly. Further design geometry modifications could be implemented to reduce high temperature exposure time for the core material. However, it seems the most feasible strategy for cell inclusion at this stage would be through post-process injection of cells (cell type selected based on application) in solution into hollow printed fibres and structures.

Figure 8. Elastic Modulus of polycaprolactone, 2% Alginate and 2% Alginate-PCL cofibres, obtained through mechanical tensile testing. The values for 2% alginate hydrogels were adapted from Drury et al [35].

a modulus of 42 ± 9 MPa and an elongation at break of 34 ± 22%. The cofibres produced are superior in strength and modulus when compared with alginate hydrogels. The PCL is observed to contribute to most of the strength of the fibre. As only 60% of the cross-sectional area of the fibre is comprised of PCL, and due to displacement of the alginate through compression from the grips of the apparatus, the modulus is reduced when compared with pristine PCL. 3.7. Cell study

A simple cell study was carried out to observe the effects of the applied shear stress and temperature of the CMEP process 7

R Cornock et al

Biofabrication 6 (2014) 025002


(B )

(C ) Figure 9. Confocal microscopy of live/dead stained cell-laden 2% alginate solutions for (A) cast-gel reference, (B) manually injected into printed hollow PCL fibres and (C) coaxially extruded cell-loaded 2% alginate-PCL composite. The fluorescent green and red observed indicate living cell metabolic activity and the stained nuclei of cells of compromised membrane integrity respectively.

4. Conclusion


The coaxial melt extrusion printing process was designed, developed and subjected to initial refinement and observation to enable the fabrication of simple, multi-layered, biomaterial constructs that combined materials with complementary properties. This system is quickly adaptable to suit the ideal processing of a range of interchangeable melt and high viscosity materials, and as such provides a potential for the fabrication of other tailored polymer nanocomposites and hydrogel models. The described coaxial melt extrusion process is capable of integrating a higher degree of functionality into tissue engineering (TE) bioscaffold technology. While many current SFF biofabrication methods lay a thermoplastic scaffold and section areas for hydrogel/cell infusion, the coaxial method allows the post-process insertion of cells and secondary material into the biomaterial fibres themselves. Coaxial scaffolds can then be infused much in the same way as their single material counterparts. This added functionality could have a significant impact on drug release platforms and TE technologies. Furthermore, the printed core-shell structures produced open avenues for exploration with respect to pathway based technology such as conducting pathways for biocompatible electronics applications, selective cell stimulation and biocompatible fluidics systems which utilize the hollow printed fibres.

The authors acknowledge the Australian Research Council Centre of Excellence for Electromaterials Science (ACES) for funding this research, the Australian National Fabrication Facility (ANFF) for provision of materials and equipment access, the use of facilities and assistance of Mr Tony Romeo at the UOW Electron Microscopy Centre, as well as the contributions of Mr Fletcher Thompson to this research. References [1] Rattanakit P, Moulton S E, Santiago K S, Liawruangrath S and Wallace G G 2012 Extrusion printed polymer structures: a facile and versatile approach to tailored drug delivery platforms Int. J. Pharm. 422 254–63 [2] Berthiaume F and Morgan J 2010 Methods in Bioengineering: 3D Tissue Engineering (Boston, MA: Artech House) [3] Sepideh H, Nsair A, Beygui R E and Shemin R J 2010 The future of cell theory and tissue engineering in cardiovascular disease: the new era of biological therapeutics Tissue Engineering ed D Eberli (Rijeka: InTech) [4] Harley B A and Yannis I V 2007 Principles of Tissue Engineering (New York: Academic) chapter 16, pp 219–36 [5] Zamora D O, Natesan S and Christy R J 2012 Constructing a collagen hydrogel for the delivery of stem cell-loaded chitosan microspheres J. Visualized Exp. 64 3624 [6] Liu X and Ma P X 2004 Polymeric scaffolds for bone tissue engineering Ann. Biomed. Eng. 32 477–86 [7] Li Z and Guan J 2011 Hydrogels for cardiac tissue engineering Polymers 3 740–61 8

R Cornock et al

Biofabrication 6 (2014) 025002

[8] Gun’ko V M, Mikhalovska L I, Savina I N, Shevchenko R V, James S L, Tomlins P E and Mikhalovsky S V 2010 Characterisation and performance of hydrogel tissue scaffolds Soft Matter 6 5351–8 [9] Mouthuy P and Ye H 2011 Comprehensive Biotechnology 2nd edn (Amsterdam: Elsevier) pp 23–36 (chapter Biomaterials: Electrospinning) [10] Lanza R P and Vacanti J P 2007 Principles of Tissue Engineering (Tissue Engineering Intelligence Unit Series) (New York: Academic) [11] Langer R and Tirrell D A 2004 Designing materials for biology and medicine Nature 428 487–92 [12] Muschler G F, Nakamoto C and Griffith L G 2004 Engineering principles of clinical cell-based tissue engineering J. Bone Joint Surg. 86 1541–58 (PMID: 15252108) [13] Hollister S J et al 2005 Engineering craniofacial scaffolds Orthod. Craniofac. Res. 8 162–73 [14] Shim J H, Kim J Y, Park M, Park J and Cho D W 2011 Development of a hybrid scaffold with synthetic biomaterials and hydrogel using solid freeform fabrication technology Biofabrication 3 034102 [15] Xu F L, Li Y B, Han J and Lv Y G 2005 Biodegradable porous nano-hydroxyapatite/alginate scaffold Mater. Sci. Forum 486–487 189–92 [16] Augst A D, Kong H J and Mooney D J 2006 Alginate hydrogels as biomaterials Macromol. Biosci. 6 623–33 [17] Wang Y, Xiao B, Dalai S, Cao X and Wu Q 2008 Development of polycaprolactone/chitosan blend porous scaffolds J. Mater. Sci., Mater. Med. 20 719–24 [18] Van de Velde K and Kiekens P 2002 Biopolymers: overview of several properties and consequences on their applications Polym. Test. 21 433–42 [19] Ikada Y and Hideto T 2000 Biodegradable polyesters for medical and ecological applications Macromol. Rapid Commun. 21 117–32 [20] Sinha V R, Bansal K, Kaushik R, Kumria R and Trehan A 2004 Poly--caprolactone microspheres and nanospheres: an overview Int. J. Pharm. 278 1–23 [21] Ben-Shabat S, Abuganima E, Raziel A and Domb A J 2003 Biodegradable polycaprolactonepolyanhydrides blends J. Polym. Sci. A 41 3781–7 [22] Xie W and Zhihua G 2009 Thermal degradation of star-shaped poly(-caprolactone) Polym. Degrad. Stab. 94 1040–6

[23] Labet M and Thielmans W 2009 Synthesis of polycaprolactone: a review Chem. Soc. Rev. 38 3484 [24] Averous L, Moro L, Dole P and Fringant C 1999 Properties of thermoplastic blends: starchpolycaprolactone Polymer 41 4157–67 [25] Rai B, Teoh S H, Hutmacher D W, Cao T and Ho K H 2005 Novel pcl-based honeycomb scaffolds as drug delivery systems for rhbmp-2 Biomaterials 26 3739–48 [26] Lam C X F, Hutmacher D W, Schantz J, Woodruff M A and Teoh S H 2009 Evaluation of polycaprolactone scaffold degradation for 6 months in vitro and in vivo J. Biomed. Mater. Res. A 90A 906–19 [27] Ms Estells J, Viduarre A, Meseguer Dueas J M and Castilla Cortzar I 2007 Physical characterization of polycaprolactone scaffolds J. Mater. Sci., Mater. Med. 19 189–95 [28] Zhou Y, Hutmacher D W, Varawan S and Lim T M 2007 In vitro bone engineering based on polycaprolactone and polycaprolactonetricalcium phosphate composites Polym. Int. 56 333–42 [29] Sharaf B, Faris C B, Abukawa H, Susarla S M, Vacanti J P, Kaban L B and Troulis M J 2012 Three-dimensionally printed polycaprolactone and beta-tricalcium phosphate scaffolds for bone tissue engineering: an in vitro study J. Oral Maxillofac. Surg. 70 647–56 [30] Schuurman W, Khristov V, Pot M, van Weeren P R, Dhert W J A and Malda J 2011 Bioprinting of hybrid tissue constructs with tailorable mechanical properties Biofabrication 3 021001 [31] Murphy M B and Mikos A G 2007 Principles of Tissue Engineering (New York: Academic) pp 309–20, chapter 22 [32] Lannutti J, Reneker D, Ma T, Tomasko D and Farson D 2007 Electrospinning for tissue engineering scaffolds Mater. Sci. Eng. C 27 504–9 [33] Cornock R, Beirne S and Wallace G G 2013 Development of a coaxial melt extrusion printing process for specialised composite bioscaffold fabrication IEEE/ASME Int. Conf. on Advanced Intelligent Mechatronics (AIM) pp 973–8 [34] Yang S, Leong K, Du Z and Chua C 2001 The design of scaffolds for use in tissue engineering: part 1. Traditional factors Tissue Eng. 7 679–89 [35] Drury J L, Dennis R G and Mooney D J 2004 The tensile properties of alginate hydrogels Biomaterials 25 3187–99


Coaxial additive manufacture of biomaterial composite scaffolds for tissue engineering.

An inherent difficulty associated with the application of suitable bioscaffolds for tissue engineering is the incorporation of adequate mechanical cha...
1MB Sizes 0 Downloads 4 Views