Introduction

The main challenge in developing scaffolds for tissue regeneration is that it is not enough for them to have the necessary biodegradability and bioactivity that stimulates regeneration [1, 2], but their mechanical and structural properties must mimic that of the native tissue, in order to activate the surrounding cells (either in vitro or in vivo). Specifically, in order for materials to be used in tissue engineering they should fulfill the following criteria: (i) they must be biocompatible, meaning that the scaffold itself should not initiate inflammation or toxicity when placed in the body, (ii) their surface should promote cell adhesion, enhance cell growth and maintain differentiated cell functions, (iii) their mechanical properties must be such that will provide the necessary structural stability while retaining flexibility and avoiding fracture upon the development of internal stresses that occur in vivo, (iv) their porosity must mimic that of the extracellular matrix such that tissue formation occurs in a uniform manner, and the pore degree/size, geometry, interconnection, and orientation must be such so they allow for the appropriate cell-cell interactions, cellular density and the necessary transfer of oxygen, nutrients and waste products [3] (it has been shown that the optimum scaffold porosity should be over 80% [4-6]), (v) depending on the application their biodegradability must be tuned.

In order for a scaffold to combine all these properties it must be a composite of materials with complementary properties, whose combination at appropriate volume fractions will allow to meet a wider range of the necessary key characteristics. In doing so the most optimum structure is the reinforcement of a soft, biodegradable polymer material, with stiff secondary phases. Such structures are the most promising for meeting both the physiological and mechanical properties of the native tissue whose regeneration is targeted.

The main polymers that are used in biomedical engineering belong to the poly (a-hydroxyester) family (poly (lactic-acid), poly (glycolic-acid) and poly(3- caprolactone)), as they are biocompatible and it is easier to mold them into various shapes and sizes. Also it is easy to tune their porosity. Numerous fabrication techniques have been developed that allow for appropriate 3D configurations to be obtained, such as 3D printing, electrospinning, hydrogels via freeze-gelation, and fused deposition modeling. These polymers have predictable mechanical properties and degradation rates, which occur through ester hydrolysis and decarboxylation from the end chains; the kinetics of which depend on the particular intermolecular interactions [7]. Other natural biopolymers such as gelatin, which is derived from collagen, and chitosan, which is a derivative from chitin, are also good candidates. These can also be considered as protein based hydrogels, but they are not the only type of hydrogels that can be used in tissue engineering.

A straightforward approach that can be followed to increase the strength of polymers is to increase the concentration of the crosslinking agent or polymer, however, it has been shown that doing so affects the diffusion of the nutrients through the scaffold and hence reduces its biocompatibility [8, 9]. Hence, in order to increase the strength of these biodegradable polymers, which is particularly necessary when reconstructing large bone segments, the addition of fibers/tubes is becoming a common technique, resulting in polymer composite structures. The “filler” materials that will be elaborated in this chapter are: boron-nitride nanotubes, magnesium alloy fibers, hydroxyapatite (HA) nanofibers, and polymer fibers. It should be noted that HA on its own is a common material considered for scaffolds and implants since it has a very high compatibility and bioactivity as it is naturally present in bone, however, its practical applications are limited due its brittleness.

 
Source
< Prev   CONTENTS   Source   Next >