Injectable scaffolds for tissue regeneration

Injectable scaffolds for tissue regeneration

(Parte 1 de 4)

Journal ofMaterialsChemistrywww .rsc .or g/mat erials

Injectable scaffolds for tissue regeneration

Qingpu Hou, Paul A. De Bank and Kevin M. Shakesheff*

Tissue Engineering Group, School of Pharmacy, University of Nottingham, University Park, Nottingham, UK NG7 2RD. E-mail:

Received 2nd February 2004, Accepted 6th May 2004 First published as an Advance Article on the web 25th May 2004

Tissue engineering aims to develop functional substitutes for damaged or diseased tissues through complex constructs of living cells, bioactive molecules and three-dimensional porous scaffolds, which support cell attachment, proliferation and differentiation. Such constructs can be formed either by seeding cells within a pre-formed scaffold or through injection of a solidifiable precursor and cell mixture to the defective tissue. As cell and bioactive molecule carriers, injectable scaffolds are appealing, particularly from the clinical point of view, because they offer the possibility of homogeneously distributing cells and molecular signals throughout the scaffold and can be injected directly into cavities, even of irregular shape and size, in a minimally invasive manner. In this paper the challenges in designing an injectable scaffold from the viewpoint of materials chemistry and the solidification mechanisms of injectable precursors are discussed. The applications of injectable scaffolds in angiogenesis, bone repair and cartilage regeneration are described.

1. Introduction

The field of tissue engineering holds great promise for the repair or regeneration of damaged and diseased tissues.1 Its underlying objective is to direct a population of cells into forming a living tissue, structurally and functionally indistinguishable from that found in nature. To guide the transition from cells to tissue, a three-dimensional scaffold may be utilized. This lends not only structural form to the cell mass, but can positively influence cell adhesion, growth and differentiation by the incorporation of adhesion molecules or the controlled release of bioactive molecules from the scaffold. As cells proliferate, deposition of extracellular matrix components and biodegradation of the scaffold results in a solid, three-dimensional tissue construct. The seeding of cells into such scaffolds can be performed in two distinct ways. Firstly, cells can be expanded in culture, seeded onto the scaffold and allowed to mature in vitro before implantation into the patient. Alternatively, the scaffold can be implanted to fill a void in damaged tissue and subsequently seeded by the infiltration of the patient’s own cells. For the latter strategy, the scaffold can either be a pre-formed, three-dimensional porous structure or an injectable scaffold; a mixture of bioactive molecules and solidifiable precursors, which are injected into the defect and form a three-dimensional structure in situ. This review will examine the benefits of injectable scaffolds, the materials chemistry challenges faced in their design, the

Qingpu Hou received his Bachelor (1989) and Master (1992) degrees from Tianjin University, China. He attained his Ph.D from the Institute of Chemistry, the Chinese Academy of Sciences in 1995. He is now a research fellow in the School of Pharmacy at the University of Nottingham, working on polymer synthesis, modification and processing for tissue engineering applications.

Paul De Bank is a postdoctoral research fellow in the School of Pharmacy at the University of Nottingham. Since completing his PhD in 2000 on the synthesis and evaluation of novel potential cannabinoids, he has worked on the spatially controlled formation of neuromuscular junctions, nonviral gene delivery systems for the tissue engineering of bone and is currently investigating novel systems for the generation of functional multicellular organoids.

Kevin Shakesheff is Professor of Tissue Engineering and Drug Delivery at the School of Pharmacy, The University of Nottingham. He became interested in tissue engineering during his time as a NATO Postdoctoral Fellow at MIT in the mid-1990s. His research focuses on the role of 3D culture environment on the behaviour of cells and the engineering of polymers to create biomimetic environments for tissue regeneration.

: 10.1039/b401791a Qingpu Hou

Paul De Bank

solidification mechanisms employed and their applications in tissue engineering.

2. The clinical need for injectable scaffolds

From a clinical perspective, the use of injectable scaffolds is very attractive as it minimizes patient discomfort, risk of infection, scar formation, and the cost of treatment. In the case of the preformed scaffolds, prior knowledge of the size and shape of the defect or cavity to be filled is necessary, and defects with irregular shape and size can prove problematical. In addition, invasive surgery for implantation of the construct is required. Moreover, cell seeding methods can be inefficient due to poor transport of cells through the matrix and cellular damage.2 The use of injectable scaffolds can overcome these limitations. By virtue of the scaffold components being in suspension or solution before solidification in vivo, a more homogeneous distribution of bioactive molecules within the matrices can readily be obtained. What is more, the nature of these systems makes it possible to co-inject a cell suspension with the scaffold components, resulting in a cell–scaffold construct that can fill any size or shape of cavity with minimally invasive surgery.3–5 After injection and solidification an in situ forming scaffold provides a temporary 3-D matrix on which the cells can adhere, proliferate and differentiate, forming a new, functional tissue.6 Hence, injectable scaffolds are promising matrices for tissue induction or regeneration, especially for engineering bone and soft tissues. In addition to serving as carriers for bioactive molecules, injectable scaffolds can also act as conduits for the guidance of tissue regeneration, tissue adhesives for healing and injectable controlled release devices for local drug delivery.7,8 Table 1 summarizes the injectable materials that have, to date, been utilized for tissue engineering applications. Prior to injection, they may be in the form of solution, paste, micro or nanoparticles, beads, or thread-like material4 and can be cell-free systems or cells and/or tissue growth factors suspension systems.

3. Materials chemistry challenges in designing injectable scaffolds

Cell viability and function within an injectable scaffold are closely related to the physical, chemical and biological characteristics of the scaffold used. From the viewpoint of materials chemistry, several requirements must be met during the design and fabrication of such a scaffold, including:

. Nontoxic and sterile components, . Injectability, . Solidification under mild conditions and cohesivity, . Mechanical strength and resistance to in situ forces, . Biodegradation, . Pore morphology, . Incorporation of bioactive molecules. These factors will be considered individually.

Nontoxic and sterile components

Injectable scaffolds should not be deleterious to the health of both cells and tissue. Each component of the formulation, the

Table 1 Injectable scaffolds reported for tissue regenerationa Injectable scaffolds Solidification mechanism References

Inorganic materials Calcium phosphate Ceramics setting 9–16 Natural polymers Chitosan Thermal gelation 17,18

Methylcellulose Thermal gelation 5 Alginate Photo cross-linking 19 Alginate Ionic gelation 20–23 Hyaluronic acid Photo cross-linking 19,24,25 Agarose Thermal gelation 26,27 Fibrin Thermal gelation 28–30 Gelatin Thermal gelation 31

Synthetic polymers Poly(aldehyde guluronate) Chemical cross-linking 32

PEG or PEO Photo cross-linking 3–42 PEO-PPO-PEO Thermal gelation 43,4 PEO-PLLA-PEO Photo cross-linking 45 PLA-g-PVA Photo cross-linking 45 PEO-PLLA Thermal gelation 46 PLGA-PEG Thermal gelation 47 PLLA-PEG Photo cross-linking 35 PEG-co-Poly(a-Hydroxy Acid) Photo cross-linking 48 PVA Photo cross-linking 49 PLAL-ASP Photo cross-linking 50 P(CL/TMC) Photo cross-linking 51–53 PLA(Glc-Ser) Photo cross-linking 54 Polyanhydrides Photo cross-linking 5,56 PPF Photo cross-linking or radical polymerization 57–65 OPF Photo cross-linking or radical polymerization 6–68 P(PF-co-EG) Photo cross-linking or radical polymerization 48,62,69–76 PhosPEG-dMA Photo polymerization 7 PNIPAAm -PEG Thermal gelation 78 PNIPAAm-gelatin Thermal gelation 31 P(NIPAAm-AAc) Thermal gelation 79,80 PEG based hydrogels Enzymatic cross-linking 81 PEG based hydrogels Michael-type addition reaction 82–84 PLA-PEG-biotin Self-assembly 2 a Abbreviations: OPF: Oligo(poly(ethylene glycol) fumarate); P(CL/TMC): Poly(-caprolactone-co-trimethylene carbonate); PDLLA: Poly(D,L-lactide); PEG: Poly(ethylene glycol); PEO: Poly(ethylene oxide); PEO-PPO-PEO: Polyethylene oxide-polypropylene oxide-polyethylene oxide; PhosPEG-dMA: Poly(ethylene glycol) di[ethylphosphatidyl(ethylene glycol)methacrylate]; PLA(Glc-Ser): Poly(L-lactic acid-co-glycolic acid-co-L-serine); PLA-PEG: Poly(lactic acid)-poly(ethylene glycol); PLAL-ASP: Poly(lactic acid-co-lysine)-poly(aspartic acid); PLGA: Poly (DL- lactic-co-glycolic acid); PLLA: Poly(L-lactic acid); PLLA-PEG: Poly(L-lactide-ethylene glycol); PNIPAAm: Poly(N-isopropylacrylamide); P(NIPAAm-AAc): Poly(N-isopropylacrylamide-acrylic acid); PPF: Poly(propylene fumarate); P(PF-co-EG): Poly(propylene furmarate-coethylene glycol); PVA: Poly(vinyl alcohol).

resulting scaffold and their degradation products should not elicit an inflammatory response or demonstrate extreme immunogenicity or cytotoxicity. Effects of the macromonomers, cross-linking agents, functional groups, initiator and residual leachable byproducts in the injection medium as well as the degradation products of the networks on cell viability and function within the scaffold should be considered.68,72,85,86 To prevent infection, the injectable precursors should be sterilizable before injection into the site of a defect, with the sterilization process having no significant impact on the chemical properties of the resulting scaffold.47,87


The precursor or macromonomer formulations should be injectable before solidification. The injectability of a scaffold is generally related to the rheological properties of the formulations, and the solidification rate or setting time of the precursors is determined by the structure/composition of the formulations and their processing conditions. For example, it has been shown that several factors including cement powder composition, liquid to powder ratio, accelerator concentration and the ageing time of the cement powder played an important role in the injectability of calcium phosphate bone cements.8 In the case of chemically cross-linked scaffolds, the chemical composition and concentration of the macromonomers and the initiators in addition to cross-linking parameters, such as the molecular weight of the cross-linking agent, are the most important factors influencing the injectability of the scaffold.39,40,68,89,90

Solidification under mild conditions and cohesivity

In a delivery system containing cells and bioactive molecules, the precursors of injectable scaffolds should undergo a mild solidification process, preferably under, or close to, physiological conditions in order to keep high cell viability and molecular bioactivity as well as avoiding damage to the surrounding tissues. Additionally, in situations where there is no defined boundary to the site of delivery, the scaffold must also be cohesive in order to enable accurate positioning at the required location. Toxic organic solvents or harsh processing conditions such as high temperature should be avoided. Ideally, the solvent used should be physiological saline, cell culture medium or a biologically acceptable organic solvent. The temperature or pH at the site of implantation should not be significantly altered during the solidification process. To this end, the solidification mechanism, the formulations of the in situ gelling precursors, and solidification process should be carefully considered.40,91 The set time or solidification time also needs to be balanced. While it should be short enough to prevent a heterogeneous distribution of cells and bioactive molecules and keep the injected gel highly cohesive at the injection site, it should also be of sufficient duration to allow for a proper surgical procedure.

Mechanical strength and resistance to in situ forces

The resultant scaffolds should have good dimensional stability in the body after injection and solidification, and possess sufficient mechanical strength to withstand biomechanical loading and provide temporary support for the cells. Additionally the mechanical properties of the scaffolds may even have an important effect on the adhesion and gene expression of the cells92,93 and should match those of the tissue at the site of implantation.94 The mechanical properties of the scaffolds largely depend on their chemical structure, crosslinking mechanism, cross-linking density, and swelling capacity under physiological conditions.25,32,49

Degradation properties

Physiological degradability of cell scaffolds is an important issue for tissue engineering applications. During tissue development, the scaffolds should degrade at a rate in line with that of new tissue regeneration. A slow degradation rate can inhibit new tissue function, as scaffold degradation can influence the deposition and distribution of extracellular matrix molecules as well as cell migration.95,96

Dissolution, hydrolysis, and enzymatic cleavage96 are three basic degradation mechanisms of biodegradable scaffolds. The main factors determining the degradation rate are the nature of cross-linking, cross-linking density, molecular weight, morphology, porosity, and amount of residual monomer. Other factors, such as the local pH and incorporation of filler, also play a role.32,65,86,97 Living cells can change the pH of their direct environment15 and consequently affect the degradation rate of the scaffold. Although a great deal of research has been performed in this area,32,54,64,73,7,98–100 the challenge still lies in the creation of injectable scaffolds with controlled degradation profiles for specific tissue regeneration.

Pore morphology

To support cell ingrowth and facilitate the exchange of nutrients and cellular waste products, the in situ formed injectable scaffolds should possess highly porous networks with specified pore morphology. In this respect, the most important parameters include pore size, porosity and interconnectivity of the porous network. To fabricate porous injectable scaffolds, several methods have been developed including in situ pore formation,101 particulate leaching, gas forming or air entraining,73,102 and automization.103

Incorporation of bioactive molecules

As most injectable scaffolds lack cell- or tissue-specific function, they should have the capacity to incorporate growth factors and cell adhesion ligands to selectively stimulate cell function and guide tissue growth. In the case of growth factor incorporation, the injectable scaffold acts as a reservoir that releases the molecules at the repair site for the length of time necessary to create an environment conducive to tissue regeneration.104 Due to the short half-lives of growth factors, an important consideration is how to retain their bioactivity and effectively release them to the site with optimal dosage, timing and, when more than one is incorporated, in the correct order.

Incorporating cell adhesion ligands, for example arginineglycine-aspartic acid (RGD)-containing sequences, into the injectable scaffolds is another approach to promoting cell specific function.105 An important issue in this respect is to select suitable peptide sequences and optimize both the density and distribution of such molecules on the scaffold surface for specific cell functions.8,3,37,98,106

4. Solidification mechanisms

Injectable scaffolds generally necessitate a means of solidifying their constituent precursors or macromonomers into a 3-D matrix in the presence of living cells within a short period of time. The solidification mechanisms directly affect the kinetics of the process and the stability of the resultant scaffolds. Typical solidification mechanisms during the scaffold formation include ceramics setting, thermally or photochemically activated radical polymerization or cross-linking,42,73,91,107 thermal gelation,108 ionic cross-linking,21,23 Michael-type addition reactions, and self-assembly mechanisms.2

Ceramics setting

Calcium phosphate cements (CPCs) can undergo a self-setting process within the body after injection, based upon the cementing action of acidic and basic calcium phosphate compounds on wetting with an aqueous medium.10 Within a few minutes, mixing of the cement formulation leads to a solidifying mass due to crystallization of dahllite.9,1 The setting time can be adjusted by addition of manipulator compounds to the wetting medium. Recently, a fully injectable calcium phosphate cement formulation was developed by incorporation of a biocompatible gelling agent. The resultant CPC had significantly improved injectability and cohesive properties.

Thermally or photochemically activated radical polymerization or cross-linking

Precursors with unsaturated or photosensitive functional groups can form gels by thermally or photochemically activated radical polymerization or cross-linking.19,40,42,73,109,110,1 Radicals produced by an initiator or photoinitiator react with the functional groups of the macromonomers to cause polymerization or cross-linking, leading to gel formation. In tissue engineering applications, the most commonly used macromonomer functional groups are (meth)acryloyl,19,35,42,45,48,50,54,107,112 styryl,112 coumarin,51,52 phenylazide,53 and fumaryl.6 Fig. 1 illustrates the chemical structures of some macromonomers that can be polymerized or cross-linked by this mechanism.107,113 The solidification process is determined by a number of factors including reactivity, functionality, concentration and molecular weight of the precursors, intensity of visible or UV light, temperature, reaction time, and the type and concentration of the initiator.

Thermal gelation

Some polymer solutions undergo gelation triggered by a change in temperature. Typical thermal gelling polymers include copolymers of N-isopropylacrylamide, poly(ethylene glycol) (PEG)-based amphiphilic block copolymers, gelatin, agarose, and cellulose.5 As one of the most intensively investigated thermosensitive polymers, N-isopropylacrylamide-based copolymers exhibit a sol–gel transition as the temperature is increased above their lower critical solution temperature (LCST) due to the drastic solubility difference of these polymers below and above the LCST. The gelation is related to the chain entanglement and the gradual chain collapse as the temperature increases.114 At room temperature, the polymer solutions are transparent and remain fluid, while at 37 uC, the matrices become opaque and form gels without significant gel induction time. The transition temperature can be further tuned by changing the composition of the copolymers. Once the gels are formed, they do not change their water content and the gelation is reversible without appreciable hysteresis. The factors determining the gelation process include polymer concentration, molecular weight, and chemical structure of the copolymer.

Recently block and star copolymers of poly(ethylene glycol) and poly(N-isopropylacrylamide) of various architectures were synthesized and their gelation behaviour studied. At low temperature, they form liquid aqueous solutions with low to moderate injection viscosities, but form relatively strong elastic gels upon warming to physiological temperature. It is believed that the linear copolymer formed a weaker gel by micellar aggregation, while the star-shaped copolymers formed a strong network gel via a physical cross-linking mechanism. This gelation is rapid and shows a low degree of syneresis.78

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