Elsevier

Biomaterials

Volume 27, Issue 35, December 2006, Pages 5909-5917
Biomaterials

The effect of anisotropic architecture on cell and tissue infiltration into tissue engineering scaffolds

https://doi.org/10.1016/j.biomaterials.2006.08.010Get rights and content

Abstract

A common phenomenon in tissue engineering is rapid tissue formation on the outer edge of the scaffold which restricts cell penetration and nutrient exchange to the scaffold centre, resulting in a necrotic core. To address this problem, we generated scaffolds with both random and anisotropic open porous architectures to enhance cell and subsequent tissue infiltration throughout the scaffold for applications in bone and cartilage engineering. Hydroxyapatite (HA) and poly(d,l-lactic acid) (PdlLA) scaffolds with random open porosity were manufactured, using modified slip-casting and by supercritical fluid processing respectively, and subsequently characterised. An array of porous aligned channels (400 μm) was incorporated into both scaffold types and cell (human osteoblast sarcoma, for HA scaffolds; ovine meniscal fibrochondrocytes, for PdlLA scaffolds) and tissue infiltration into these modified scaffolds was assessed in vitro (cell penetration) and in vivo (tissue infiltration; HA scaffolds only). Scaffolds were shown to have an extensive random, open porous structure with an average porosity of 85%. Enhanced cell and tissue penetration was observed both in vitro and in vivo demonstrating that scaffold design alone can influence cell and tissue infiltration into the centre of tissue engineering scaffolds.

Introduction

The clinical need to develop improved material structures for reconstructive surgery and tissue engineering has been the driving force behind both academic and commercial research into regenerative medicine [1], [2], [3]. Many tissue engineering strategies involve the fabrication of a material scaffold that plays an important role in supporting initial cell attachment and subsequently in the guidance of tissue formation whilst providing mechanical stability. These devices can be produced from synthetic and naturally derived polymers, as well as inorganic materials, and are widely used in the field of tissue engineering [4], [5], [6].

Much research within tissue engineering focuses on the design of the scaffold to ensure appropriate tissue growth is encouraged [7]. However, a common problem encountered when using such scaffolds for tissue engineering is the rapid formation of tissue on the outer edge, which leads to the development of a necrotic core due to the limitations of cell penetration and nutrient exchange [8], [9]. Most tissues possess a network of blood and capillary vessels that perform this function in vivo but engineering such a complex construct in vitro is challenging and as yet, has eluded tissue engineers [10], [11]. Even those tissues that are avascular, such as cartilage, can prove challenging to engineer if the in vivo nutrient supply cannot be adequately simulated in vitro [8].

To date many different strategies have been employed to overcome this issue. A common approach is to utilise sophisticated culture systems to perfuse culture media around and/or through the scaffold [12], [13], [14]. Although these bioreactors have been successful and are increasingly vital to tissue engineering in vitro, this field is still developing appropriate systems for good tissue growth. However, even in an optimised in vitro culture system, there is still a need to ensure tissue growth occurs evenly throughout the construct. Furthermore, for strategies where the scaffold alone will be implanted and the body used as the bioreactor, it is important to ensure that native tissue can infiltrate the whole scaffold to ensure adequate integration of the construct [15]. Another strategy to encourage tissue formation and cell differentiation within scaffolds is to incorporate growth factors which can also act as chemotactic agents to encourage cell migration [3]. Enhanced tissue formation using this method has been demonstrated with the PdlLA scaffolds presented in this study [16] but there is still scope to improve this further by potentially using a dual strategy.

An alternative method of addressing these issues is to incorporate a design within the scaffold that will improve nutrient and cell transfer to the scaffold centre, both in vitro and in vivo. This is especially important whilst the challenge of incorporating a vasculature into tissue engineered constructs (that require a blood supply) prior to implantation is being resolved to ensure necrosis does not occur within the scaffold during the establishment of a native vasculature. Recent studies have reported the incorporation of aligned channels into the general structure of the scaffold in order to achieve this goal. This has been demonstrated using fabrication methods such as phase-separation techniques [17], solid free-form fabrication [18] and rapid prototyping [19]. The resulting scaffolds with aligned channels essentially provided a thoroughfare for both cell and nutrient flow. In all these studies, where either in vitro and/or in vivo studies were carried out, cell/tissue penetration into the scaffold was observed. However, it is worth noting that some of these studies still reported greater cell/tissue infiltration at the periphery of the scaffold rather than in the centre and a comparison to nonarchitectured scaffolds was not made [19], [20], [21].

Previous studies from our laboratories have demonstrated that the optimum channel diameter for human osteosarcoma cell penetration into porous hydroxyapatite (HA) scaffolds was approximately 400 μm [22]. This work is taken further in this paper where the strategy of incorporating anisotropic macroarchitecture within porous scaffolds is assessed both in vitro and in vivo in two scaffold types suitable for bone and cartilage tissue engineering.

Section snippets

Ceramic hydroxyapatite scaffolds

HA scaffolds were manufactured using a modified slip-casting method as described previously [22], [23]. For scaffolds with aligned macroarchitecture, stainless steel needles (600 μm diameter) were inserted, at the green state, perpendicular to the cylindrical face (1 centrally located channel for in vitro studies and 13 channels for mechanical testing and in vivo studies). After sintering in air at 1350 °C for 3 h, the ceramics for biological studies were washed by ultrasonic agitation in ultra

Scaffold characterisation

Using a modified slip-casting methodology and scCO2, 3D random porous and interconnected HA scaffolds (Fig. 1A and B; 10 mm diameter×5 mm height for in vitro studies and 5 mm diameter×5 mm height for in vivo studies) and PdlLA scaffolds (10 mm diameter×5 mm height; Fig. 1C and D), respectively, were fabricated. HA scaffolds were on average 85% porous with an average pore size of 594±599 μm (±SD; Fig. 2A) and an interconnecting window size of 368 μm with a range of 16–1400 μm (Fig. 2B). Helium pycnometry

Discussion

Scaffold design is becoming increasingly important for tissue engineering applications as scaffold morphology, porosity, pore size and degree of interconnectivity are all considered important for cell penetration and subsequent tissue infiltration and development. We recently reported the fabrication of random, porous HA scaffolds for bone tissue engineering, using a modified slip-casting method [23] and have extensive experience in the fabrication of PdlLA scaffolds using scCO2 technology [24]

Conclusion

In summary, we have fabricated both ceramic and biodegradable polymer tissue engineering scaffolds with a random, open porous architecture with and without aligned channels. We have characterised these scaffolds in terms of pore and window size and mechanical strength. In vitro and in vivo data demonstrated enhanced cell penetration and tissue infiltration throughout the scaffold structure. We believe that a highly porous isotropic microstructure, combined with an anisotropic macroarchitecture,

Acknowledgments

The authors would like to acknowledge the Foresight LINK consortium, Department of Trade and Industry (DTI) and EPSRC for funding. The authors are also grateful to the orthopaedic surgeons at the Southampton General Hospital for their aid in facilitating bone marrow sample collection.

References (41)

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1

Current address: Skeletal Tissue Engineering Laboratory, Leeds Dental Institute, University of Leeds, Leeds LS2 9LU, UK.

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