Tailoring Bulk and Surface Composition of Polylactides for Application in Engineering of Skeletal Tissues

  • Gloria Gallego Ferrer
  • Andrea Liedmann
  • Marcus S. Niepel
  • Zhen-Mei Liu
  • Thomas Groth
Part of the Advances in Polymer Science book series (POLYMER, volume 282)


Synthetic biodegradable polylactides have been used extensively to fabricate scaffolds for engineering skeletal tissues such as bone and cartilage. This chapter summarizes the application of polylactides in tissue engineering and shows strategies for tailoring its bulk and surface composition for optimized degradation rates, mechanical properties, and bioactivities that cannot be achieved with pure polylactide polymers. Hence, block copolymers and the use of blending as a cost-effective strategy are described here. Furthermore, polymeric networks are shown that are advantageous in porogen-leaching manufacture of scaffolds, in preventing crystallization during degradation, and in allowing the incorporation of hydrophilic chains. In addition, mechanical reinforcement of the polymer is achieved when organic–inorganic composites of polylactides are formed. The last part of this chapter focusses on the modification of the surface to tailor the biocompatibility of polylactides only, without changing the bulk properties of the material. Surface modification by wet chemical processes and adsorption of biogenic multilayers of glycosaminoglycans is described that not only significantly improves biocompatibility but may also help to drive differentiation of stem cells into the desired lineage.


Blending Bone Cartilage Composites Copolymer Crosslinking Polyelectrolyte multilayers Polylactides Polymeric networks Surface modification 

1 Tissue Engineering of Skeletal Tissues

Tissue engineering is an interdisciplinary field that applies the principles of engineering and biomedical sciences to the development of biological substitutes that can restore, maintain, or improve tissue or organ functions that are defective or have been lost as a result of trauma or disease [1]. The rapidly developing field of regenerative medicine requires rational molecular and supramolecular design of temporary scaffold materials for cells in order to control their bioactivity and physical properties [2, 3, 4].

The ideal scaffold material is biocompatible, nontoxic to cells or the surrounding tissue, biodegradable, chemically and mechanically stable, and has low immunogenicity [5, 6]. Most biomaterials should be structurally stable for long enough to allow repaired or regenerated tissue to organize into a desired three-dimensional (3D) structure [7, 8]. The material should ideally degrade without any toxic residues remaining [9]. Furthermore, the microstructure and porosity of the material should be controllable, so that biomaterials provide structural support to the surrounding tissue or encourage cell proliferation, spreading, and differentiation [5, 10, 11]. This can be accomplished with techniques that include modifying the material’s wetting properties, charge, pore structure, surface topography, or functionalization of the material surface with extracellular matrix (ECM)-like molecules or therapeutic proteins [12, 13, 14, 15]. Ideally, a given biomaterial can be formed or processed into a variety of shapes such as tubes, sheets, meshes, sponges, and foams. In cases in which local delivery of potentially therapeutic molecules is desired, the materials should permit a controlled and sustained delivery of those molecules for the required duration [5, 16].

Tissue engineering of skeletal tissues reveals several challenges, in spite of their relatively simple structures. Skeletal tissues consist of highly organized 3D networks of cells and matrix, giving rise to tissue structures with remarkable mechanical properties. The cells continually remodel the tissue and respond to changes in the physical and mechanical environment of the body. The building of functional tissue addresses very complex biological problems, especially creation of higher order skeletal structures such as the zonal organization of articular cartilage or a complete joint with integrated cartilage and bone tissue. For roles that purely nonbiological prosthetics are incapable of fulfilling, the engineering designs may require control of the spatial organization of multiple cell types in a 3D system that encourages the production of the correct types and amounts of matrix molecules in the appropriate structural framework [17]. Therefore, the next generation of engineered musculoskeletal tissues needs to be more complex and structurally organized to mimic normal tissue structure and function more fully.

For example, slow degradation is required for tissue engineering of the skeletal system, whereas for skin the scaffold does not need to stay in place for longer than 1 month. Different requirements for a scaffold depend on the different biological properties of normal articular tissue. Bone is a self-repairing structural material, but adult articular cartilage contains no blood supply, neural network, or lymphatic drainage and rarely heals, even with continuous passive motion. Therefore, tissue engineering strategies for cartilage attempt to introduce cells to the injury site, with or without a scaffold, to take over the role of ECM production and remodeling. Bone tissue engineering efforts, on the other hand, focus on implantation of biomaterials and factors that augment the natural repair process in larger defects [17]. A considerable challenge in skeletal tissue repair is the fact that the scaffold must be able to withstand physiologic loading until sufficient tissue regeneration occurs. Most soft porous biomaterials such as collagen (COL) are not able to endure fixation with sutures because of their low tearing strength. In such cases, reinforcement (e.g., with resorbable fibers) is required [18].

Synthetic biodegradable poly(α-hydroxy ester)s or their copolymers, such as poly(lactic acid) (PLA) have been used extensively to fabricate scaffolds for engineering skeletal tissues such as bone [19] and cartilage [11, 20, 21, 22]. They enable relatively good mechanical and manufacturing properties, moderate biocompatibility, and precise control over physiochemical properties such as degradation rate, porosity, and microstructure and over mechanical properties [23]. Nevertheless, the vast majority of biodegradable polymers appear to have obvious drawbacks that hinder their broader application.

In the next section, the advantages and disadvantages of the biodegradable polymer PLA will be discussed. Subsequent sections deal with the possibilities of modifying PLA to overcome its negative properties or adapting it to a particular application. Section 3 focusses on strategies for tailoring its bulk composition to develop functional scaffolds. Implants and tissue engineering scaffolds interact with the biological environment via their surface. Bulk and surface properties of biomaterials used for implants have been shown to directly influence and, in some cases, control the dynamic interactions that take place at the tissue–implant interface [13]. Modification of the outermost part of materials may be sufficient to tailor their biocompatibility, which extensively affects the cell response if the materials are exposed to a biosystem [24]. Based on this, the strategies for biomaterial surface modification have been adapted over the years to improve the surfaces for desired applications. Surface modification is a feasible approach for making hydrophobic biomaterials more wettable or introducing biospecific cues to promote adhesion of cells [4, 25, 26]. Section 4 describes selected strategies for surface modification of polylactides by chemical and physical modification, which can be used to increase cell adhesion and mesenchymal stem cell (MSC) differentiation.

2 Advantages and Disadvantages of Using Polylactides in Tissue Engineering

The vast majority of biodegradable polymers studied belong to the polyester family, which includes polyglycolides and polylactides. PLA belongs to a family of linear aliphatic polyesters of α-hydroxy acid derivatives. In general, PLA can be obtained either by direct condensation of lactic acid or by the ring-opening polymerization of the cyclic lactide dimer [27] (Fig. 1).
Fig. 1

Polymerization routes to obtain PLLA (DP direct polymerization, ROP ring-opening polymerization)

Because lactide has two asymmetric carbons, three optically active isomers are possible, l- and d-stereoisomers and the racemic d,l-form. These different isomers give rise to a range of different polylactides, including poly(l-lactide) (PLLA), poly(d,l-lactide) (PDLA), and poly(l-lactide-co-d,l-lactide) (PDLLA). The chirality of lactide unit provides the means to adjust the degradation rate as well as the physical and mechanical properties. PLLA and PDLA are semicrystalline solids with similar hydrolytic degradation rates as poly(glycolic acid) (PGA). However, PLLA is more hydrophobic and also more resistant to hydrolytic attack than PGA. For most applications, the l-isomer of lactic acid is usually preferred because PLLA offers the best compromise in terms of mechanical stability and degradation rate [9]. These properties favor its use for most clinical applications, as does its more optimal metabolism in vivo.

The degradation of PLLA and its copolymers generally involves random hydrolysis of their ester bonds. Hydrolysis begins with diffusion of water into the amorphous regions, and proceeds by hydrolytic cleavage, which occurs from the edge toward the center of crystalline domains. This process results in a decrease in molecular weight (MW) followed by a reduction in mechanical properties and loss of mass. PLLA degrades to lactic acid, which is generally present in the metabolic pathways of mammals. The natural pathways (metabolism, excretion) of all animals and microorganisms eliminate the final degradation products [28]. Eventually, PLLA enters the tricarboxylic acid cycle and is digested into water and carbon dioxide [29]. Administration of carbon-13 (C13)-labeled PLA results in little radioactivity in feces or urine, indicating that most of the degradation products are released through respiration [30, 31]. Because PLLA is more hydrophobic and less crystalline than PGA, it degrades at a slower rate [7]. The degradation rate of the amorphous copolymer can be easily controlled by altering the ratio of PLA to PGA in the formulation when blends or copolymers are produced.

In addition to biodegradability, an advantage of polyesters is their simple and scalable manufacture into a variety of different shapes. Polyesters such as PGA and PLA can be easily formed into desired forms by molding, protrusion, and solvent processing [32], which makes them ideal candidates for making scaffolds for tissue engineering applications. For formation of PGA- and PLA-derived scaffolds, the solvent-casting method is applied when a polymer mesh is required, which is precipitated from viscous solutions into a low toxicity laminar film. Macroporous PLA scaffolds with microporous trabeculae can be prepared by freeze-extraction and particle leaching processes using dioxane as solvent (and also microporogen) and spherical polymeric particles as macroporogen. Applying this technique, we performed a systematic variation of polymer/dioxane and polymer/porogen ratios to manufacture PLLA scaffolds with a broad range of physical properties [33]. The dioxane proportion was varied until the limit of processability. Experimental results of different uniaxial compression tests, permeability studies, porosity, and pore size distribution function were properly correlated with the synthesis parameters to quantify solvent and mass weight of porogen. The obtained structures were observed with scanning electron microscopy (SEM), as depicted in Fig. 2. Macropores ranged from approximately 108–164 μm, whereas micropores were 1.3–13.7 μm. The micrographs showed a clear decrease in micropore size with a reduction in the amount of dioxane. A higher quantity of porogen spheres produced a closer connection between them, leaving bigger pores after their removal. However, increasing the porogen concentration (from 1:1 to 1:1.25) had a slight influence on macropore size (Fig. 2).
Fig. 2

Scanning electron micrographs of PLLA scaffolds prepared from different PLLA/dioxane solutions (10, 15, and 18 wt% of PLLA ). (ac, gi) Low magnification images showing the macroporous structure for two series (1:1 and 1:1.25 PLLA solution/porogen weight ratio). (df, jl) High magnification images showing the microporosity and trabecular structure of the scaffolds. Reprinted with permission from [33]

These polymers offer distinct advantages. To date, PLLA has been used for the regeneration of different tissues and organs, such as liver [34], bone [35], cartilage [20, 22], and skin [36]. Furthermore, it is a suitable substrate for osteoblast culture and exhibits bone fracture healing capabilities, with results comparable to those obtained with metallic implants [37]. In addition, chondrocyte differentiation is significantly improved in the presence of PLLA [38]. Furthermore, the in vivo degradation rates of PLLA are similar to those exhibited in vitro [21].

On the other hand, many studies have shown that the aforementioned polymers and their copolymers do have some limitations in tissue engineering applications. Some disadvantages of these polymers are their lack of cell adhesion motives, hydrophobicity, self-acceleration of degradation (autocatalysis), release of acidic degradation products, high melting points that influence processability, and early decline of mechanical properties during degradation. These properties are serious drawbacks to their medical application [8].

PLLA is a relatively hydrophobic polymer, whose biocompatibility depends on its molecular weight and crystallinity [19]. Crystallization kinetics of PLLA from the melt is slow and depends remarkably on the cooling rate [39, 40]. Amorphous PLLA is usually obtained by cooling it from the melt at a broad range of rates. Amorphous samples can then be placed at the cold crystallization temperature for different times to produce materials with different crystallinities [41]. Crystallinity and crystal size are parameters that regulate PLLA biocompatibility and its interaction with cells [42]. However, the main importance of these parameters is that they modulate the degradation kinetics of the polymer. Polylactides undergo degradation mainly through hydrolysis, which is initiated in the amorphous phase of the polymer [43] and its rate strongly depends on the initial crystallinity of the polymer [44, 45]. Tailor-made scaffolds with desired degradation rates can be produced from amorphous PLLA scaffolds by allowing them to crystallize at the cold crystallization temperature.

The degradation products of polylactides are acidic and can be toxic to implanted cells if released in large quantities [46]. Some studies have shown systemic or local reactions caused by acidic degradation products [28]. Concerns about the biocompatibility of PLLA were heightened following the demonstration that PLLA produced toxic solutions, probably as a result of acidic degradation, usually in cases where the implant used was of considerable size [47]. Another concern is the release of small particles during degradation in vivo, which can cause an inflammatory response through phagocytosis by macrophages and multinucleated giant cells [48]. It has also been noted that no adverse biological responses occur in cases where the volume of the material used is relatively small [9]. However, the biocompatibility of polylactide-based scaffolds is not high for all type of cells due to the hydrophobic nature of the polymer [49], which may result in impaired adhesion of human mesenchymal stem cells (hMSCs). The low wettability is related to these findings [50]. It has been shown in a number of studies that hydrophobic biomaterial surfaces hamper attachment of cells and/or impair their subsequent growth and function [51]. The underlying reasons for this phenomenon are the altered adsorption and conformation of adhesive proteins such as fibronectin (FN), which may lead to impaired interaction with cellular adhesion receptors, namely integrins [52, 53]. Furthermore, the adsorption of blood cells on porous polylactide scaffolds indicates that these scaffolds are highly thrombogenic [54]. Additionally, the absence of reactive functional sites on their chains is a further limitation to their application as scaffolds. Because of these obvious drawbacks regarding control of degradation and corresponding mechanical properties, generation of acidic by-products, and limited support for cell adhesion, the modification of bulk composition and surface of polylactides has been introduced for applications in tissue engineering [55]. These modifications are described in more detail in Sects. 3 and 4.

3 Tailoring the Bulk Composition of Polylactides

Despite the benefits of pure PLLA in cartilage and bone tissue engineering [23, 42, 56, 57], there is extensive literature on tailoring its bulk composition to develop scaffolds with properties that best fit the properties of these tissues. The most common modifications are aimed at regulating the degradation rate to align it with the real production of new ECM and at achieving mechanical properties more similar to those of the specific tissue. One problem is that polylactides have long degradation rates of about 1 year or more. Even if hydrolysis first affects the amorphous phase, completely amorphous samples have longer degradation rates than semicrystalline samples, because degraded chains, instead of being released from the polymer, remain in the samples and start to crystallize [44]. This means that there is a limit to the degradation rate, which cannot be overcome by modulating the crystallinity of the sample. Faster degradation rates can be obtained by combining polylactides with other hydrophobic polymers with faster hydrolysis rates or with hydrophilic polymers that speed up water absorption. Regarding the mechanical properties, the same modifications that accelerate the degradation rate can give rise to softer materials that are desirable, for instance, in cartilage regeneration [58]. When reinforcement is needed (e.g., in bone), organic–inorganic composites can be effective for increasing the elastic modulus, which in turn improves bioactivity [59, 60]. The ECM of cartilage is highly hydrated and, consequently, hydrophilic materials such as hydrogels are preferred as a synthetic matrix to replace this tissue [58]. However, hydrogels are extremely soft and polylactides that can provide mechanical stability seem to be a better alternative. However, the cell adhesive pore walls of PLA-based scaffolds are not adequate for chondrocytes because they lose their phenotype when adhering to the pore surface [61]. Hence, polylactides can be combined with hydrophilic polymers to reduce cell adhesiveness, such as cylindrical osteochondral scaffolds based on PLA [62, 63]. Devices less invasive to the subchondral bone have been proposed for regenerating the articulation as a whole [64].

We summarize here the main strategies described in the literature for optimizing bulk materials in the search for optimized degradation rates, mechanical properties, and bioactivities that cannot be achieved with pure PLA polymers. We start by reviewing block copolymers, because the resulting morphologies are more predictable than those obtained by blending. Then, we summarize the main blends and discuss the use of blending as a cost-effective strategy. Polymeric networks are systems that are not soluble in good solvents of the polymer, but are advantageous in porogen-leaching manufacture of scaffolds, in preventing crystallization during degradation, and in allowing incorporation of hydrophilic chains. Because of their importance in regeneration of bone, we also review organic–inorganic composites as bioactive systems with good mechanical properties.

3.1 Block Copolymers

Three kinds of copolymers are usually synthesized: AB diblock, ABA triblock, and multi-arm block copolymers. Because the blocks in the copolymers are usually immiscible, they have phase-separated morphologies with ordered domains, forming a variety of nanostructures, which can also be crystalline [65]. Despite phase separation, because the blocks correspond to the same chain, the obtained domains are normally smaller than found in blends and the resulting properties are different. PLA is usually copolymerized with PGA to obtain poly(lactide-co-glycolic acid) (PLGA) copolymers, with degradation rates regulated by the ratio of their respective monomers [66, 67]. In contrast to PLA and PGA, PLGA copolymers are amorphous with glass transition temperatures (Tgs) in the general range of 45–55°C, depending on the lactide and glycolide ratio [68]. Glycolic acid units are more hydrophilic than lactic acid units, meaning that copolymers rich in the former present faster degradation than copolymers rich in lactic acid, with the exception of the 50:50 copolymer that presents the highest degradation rate [69]. The most commercialized compositions for medical products are 50:50, 65:35, 75:25, 85:15, and 90:10 [70]. Despite the advantages of these copolymers, their toughness is too high for certain biomedical applications, and copolymerization with more flexible chains is recommended. This is the case for l-lactide and ε-caprolactone copolymers (PLCL) that are more flexible than PLA and PLGA thanks to the low Tg of poly(ε-caprolactone) (PCL) (−60°C) and consequently higher mobility of caprolactone blocks at body temperature [71]. For lactic acid contents higher than 30 wt%, copolymers are amorphous with Tg values increasing with lactide content, which also affects the mechanical properties [72]. Degradation rates of these copolymers are usually faster than those of the homopolymers and can be varied over a wide time range [73]. Such elastic, biodegradable PLCL scaffolds are able to deliver effective mechanical stimulation to cells and could be useful for cartilage regeneration [74].

Amphiphilic copolymers of PLA are very interesting because they are able to reduce PLA brittleness and increase surface hydrophilicity, with controlled degradation rates. Poly(ethylene glycol) (PEG) block copolymers with PLA have been proposed for making scaffolds due to the good biocompatibility and nontoxicity of PEG. ABA triblock copolymers of PLLA-b-PEG-b-PLLA (PELA) exhibit a good balance between degradation rate and hydrophilicity [75]. The Tg depends on several factors, but is usually lower than that of pure PLA. As described [76], the Tg ranges between 15 and 35°C. The melting peak is also shifted to lower temperatures and lies in the range of 60–120°C. These copolymers form a microphase-separated structure with hard domains alternating with soft domains, forming a thermoplastic physical hydrogel [76, 77].

The combination of PLA with polysaccharides is also an interesting strategy for increasing the hydrophilicity and modulating the degradation rate of polyesters, although the main application of these copolymers is focused on drug delivery systems. For instance, amphiphilic brush-like copolymers of PLA and dextran have been proposed as compatibilizers of PLA/dextran blends, giving rise to blends with improved mechanical properties [78]. In-situ forming microgels of hyaluronic acid (HA) grafted with PLA find application in intraarticular administration of anti-inflammatory drugs [79].

3.2 Blends

Polymer blending is perhaps the most economic solution for modifying the properties of PLA. However, as most polymers are immiscible, polymer blending gives rise to phase-separated materials with domains that are usually larger than those obtained in copolymers. Furthermore, the properties of the resulting blend can be poorer than those of pure polymers if the microstructure is not adequate [80]. As recently reviewed, compatibilization strategies can effectively enhance PLA properties in blends, but care has to be taken to maintain the material’s biocompatibility for biomedical applications [81].

PLLA and PDLA of average molecular weights below 1 × 105 g mol−1 form miscible blends because of their ability to form stereocomplexed crystallites that melt at about 230°C, which is 50°C above the melting temperature of PLLA or PDLA homocrystals [82, 83]. Stereocomplexation provides higher mechanical and thermal stability to blends and extends the range of application of polylactides in medicine [84]. As a consequence of stereocomplexation, PLLA/PDLA blends show slower degradation rates than PLLA, which results in milder inflammatory reactions in vivo [85].

Literature referring to PLA/PGA blends is rare, probably because phase separation gives rise to systems with properties that are not advantageous compared with the copolymers. On the other hand, the resulting morphology of these blends allows preparation of very interesting porous fibers by selectively solubilizing the PLA phase of electrospun membranes [86]. PLLA blended with PLGA is more common and can lead to materials with tailored degradation rates and mechanical properties that depend on the molecular weight of the pristine polymers, lactide-to-glycolide ratio in PLGA, and the blending ratio of both components [83]. A systematic study on a range of compositions of PLLA and PLGA demonstrated that for a certain composition (PLA/PGA 3:1), the blend is partially miscible and presents a Young’s modulus higher than that predicted by the additive rule of phases [87]. This partial miscibility also accelerates the degradation of the PLLA matrix, which can be advantageous in tissue engineering.

Blending of PCL with PLLA increases ductility and improves the ultimate strength of PLLA [83]. Although both polymers are immiscible, dispersing PCL domains in the PLLA matrix increased elongation at break, toughness, and thermal stability of PLLA as a result of the plasticizing effect of the mixed amorphous phase and reduction in PLLA crystallinity [88]. Block copolymers of PLA and PCL can be added to blends to increase their compatibility [89], giving reduced PCL domain size when the copolymers are present.

In a similar way as for copolymers, PEG blending with PLA leads to more ductile materials with an accelerated degradation rate [83]. Blending avoids the use of toxic catalysts employed in copolymer synthesis (e.g., stannous octoate) [90]. More importantly, PEG chains in the blends act as plasticizer, facilitating PLA extrusion [91] and direct 3D printing [92].

Attempts have been made to mix PLA with other hydrophilic biomaterials of natural origin such as chitosan (CHI), starch, or gelatine (GEL) by the use of a common solvent [62, 93, 94, 95]. However, in these cases, immiscibility is more marked than when two hydrophobic polymers are mixed and processing processes need to be modified to prevent large phase-separated domains that could compromise the properties of the blends. For instance, a PLA/CHI solution was precipitated in acetone [93], which is a very fast way to eliminate the solvent, preventing phase separation to some extent. Dimethyl sulfoxide (DMSO) is a common solvent of PLA and GEL and, although phase-separated blends were obtained [95], the GEL-dispersed domains in the blends with a content below 5% were of several hundred nanometers, without reaching the micrometer scale. Furthermore, the mechanical properties did not decrease much with respect to pure PLA. GEL provided hydrophilicity and prevented PLA crystallization, and the low content GEL mixtures were better than pure materials in terms of degradation profile. As a result, hydrogel blends of PLA had increased hydrophilicity, improved biocompatibility, and did not necessarily compromise the mechanical properties of PLLA if present in a low percentage [25, 96]. A good example of this combination is an osteochondral scaffold [62] made of a blend of starch and PLLA on the cartilage side, which was found to possess adequate hydration capability. For the bone region, where more stiffness and strength is required, PLLA reinforced with hydroxyapatite (HAp) was used. PLLA elasticity was also regulated by mixing with the microbial polyester poly(3-hydroxybutyrate-co-3-hydroxyvalerate) (PHBV) [97], which enhanced cell biocompatibility, particularly for the 60:40 and 50:50 composition ratios [98].

3.3 Polymeric Networks

A polymer network is formed when a polymer is chemically crosslinked. If the crosslinker has two or more functionalities, it is able to bind to two or more polymer chains at the same time, forming a network that swells in the presence of a proper solvent but does not dissolve in it. The interest on polymeric networks of PLA in tissue engineering is dual. On the one hand, PLA allows the preparation of scaffolds using porogen particles that need to be removed with organic solvents that are also good solvents of PLA. For instance, porogens are present in templates formed by sintered polymeric particles [99]. The precursor solution of the polymer is injected in the voids of the template, polymerized and/or crosslinked, and the porogen is then removed by dissolution with an organic solvent that must not dissolve the polymer; this process generates a porous structure [99]. It is difficult to find a combination of solvents for this technique because the porogen solvent also usually dissolves the PLA. The problem can be avoided by crosslinking PLA in the presence of the porogen to avoid its dissolution when removing the porogen.

On the other hand, polymer network formation can prevent PLA crystallization, giving rise to more ductile materials. It also permits the combination of PLA with other polymers in the network, which can regulate PLA hydrophilicity, crystallinity, and degradation rate.

Methacrylate end-capped PLA macromers were successfully synthesized through condensation of PLA diol with methacrylic anhydride [100]. When a solution of this macromer is exposed to UV light in the presence of an initiator, the double bonds at the end of the chains are opened and radicals react with each other, giving rise to PLA networks. The hydrophilicity of PLA in these systems can be tuned by adding different amounts of an acrylic, hydrophilic monomer, which accelerates hydrolytic and enzymatic degradation of PLA and promotes MSC differentiation to the osteogenic phenotype [26, 100]. PLA crystallization is prevented by network formation and the Tg of the polyester domains shifts from 45°C in the pure network to 40°C or less in the hydrophilic networks.

In a similar way, polymeric networks combining PLA and PCL can be prepared by crosslinking solutions of methacrylate-terminated PLA and PCL macromers at different proportions [101]. This strategy produces quasi-compatibilized networks with a single Tg in the differential scanning calorimetry (DSC) thermograms that is lower than for the pure PLA network and splits into two relaxations in the differential mobility spectrometry (DMS) spectra, indicating segregation into small domains of pure PCL. Combined networks are amorphous, except for a 70 wt% of PCL. Network formation from macromer solutions in the voids of a polymeric template allowed synthesis of interconnected macroporous scaffolds with different hydrophilicities that promoted osteogenic differentiation of MSCs [102] and chondrogenic differentiation of de-differentiated chondrocytes [103].

PLA/CHI networks with short degradation times regulated by the amount of PLA were successfully obtained by grafting PLA chains onto CHI in the presence of stannous octoate and trimethylamine, subsequent methacrylation of the end group of grafted PLA chains, and final exposure to UV light [104]. Because PLA chains acted as crosslinker in these networks, the swelling capacity of CHI increased with the addition of PLA, thus reducing the compressive modulus. This system demonstrated injectability into the body and is a proper hydrogel scaffold for tissue engineering of bone because it efficiently released BMP-2 growth factor and induced osteoblastic differentiation and mineralization.

Hydrogel networks formed from multifunctional macromers of PLA-b-PEG-b-PLA end-capped with dimethacrylate (DM) groups (PEG-LA-DM) are also injectable hydrogels that find application in tissue engineering [105]. Their combination with a slowly degrading macromer such as PEGDM leads to networks with different degradation rates that can induce osteogenesis and mineralization, both of which increase with degradation time.

3.4 Organic–Inorganic Composites

Hard tissues such as bone require scaffolds able to sustain the load applied in vivo. Ceramic materials are brittle and stiff, and therefore macroporous materials made with them are fragile. By contrast, most polymeric materials are unable to sustain the applied loads in bone [106]. Of the biodegradable polymers, PLA offers good mechanical properties and has been proposed for tissue engineering of bone. However, in macroporous scaffolds, the mechanical properties are affected by the porosity and reinforcement is a recurrent issue in bone replacement. Nanocomposites of a PLA matrix reinforced with HAp or other ceramic material such as tricalciumphosphate (TCP), silica, or bioactive glass have been developed with the aim of enhancing the mechanical properties of PLA and designing porous materials for tissue engineering of bone [107]. The popularity of these composites lies not only in the mechanical reinforcement; the inorganic filler also acts as a buffer for the decrease in local pH caused by PLA degradation products. Hence, the filler modulates degradation rate and enhances cell adhesion, osteoconductivity, and bioactivity of the PLA scaffold [107, 108]. Bioactivity has been defined as the ability of a biomaterial to bind bone directly without forming any surrounding connective tissue. Of the bioactive inorganic fillers, HAp is preferred because it has the same formulation as mineral bone [Ca10(PO4)6(OH)2] and, from the crystallographic point of view, is the most similar to natural bone [108].

Finding the adequate amount of inorganic filler, properly dispersing it in the polymeric matrix, and enhancing the interfacial bond strength with the matrix are probably the most discussed challenges in obtaining the required bioactivity and reinforcement for these composites [109]. Sonication of a solution of PLA in dioxane containing HAp particles and subsequent freeze-extraction leads to proper dispersion of HAp in membranes [110] and scaffolds [111]. As an example, Fig. 3 shows the SEM image of a PLLA membrane containing 10 wt% of HAp prepared by freeze-extraction, together with the corresponding energy dispersive spectroscopy (EDS) scan that confirms the effective incorporation of HAp (Ca and P peaks) [110]. The scan shows that HAp particles are homogeneously dispersed in the walls of the pores.
Fig. 3

Scanning electron micrograph (SEM) and energy dispersive spectroscopy (EDS) analysis (inset) of a membrane of PLLA/HAp nanocomposite with 10 wt% of HAp and a surface plasma treatment to increase its hydrophilicity. The membrane was manufactured by freeze-extraction from a 15 wt% of PLLA solution in dioxane. Reprinted with permission from [110]

To improve bioactivity, these composite scaffolds can be mineralized in vitro by immersion in simulated body fluid (SBF) prior to implantation. Their integration in the host bone and the bone regenerative potential is superior to that of nonmineralized composites or bare PLA scaffolds, as described previously [111]. In that study, we demonstrated that PLLA scaffolds, PLLA/HAp composite scaffolds with 5 wt% of HAp, and PLLA/HAp scaffolds coated with biomimetic apatite (by immersion in SBF) (PLLA/HAp/SBF) respond differently when implanted into the bone of sheep [111]. Masson’s trichrome staining histologies of implanted scaffolds and surrounding tissue are seen in Fig. 4a, and the detail of the interface between the scaffolds and the host bone in Fig. 4b, demonstrating good integration of the three scaffolds with the host tissue and no fibrous capsule formation. The microscopic aspect of the regenerated tissue varied strongly with scaffold composition (Fig. 4c). PLLA with and without HAp was invaded by a fibrous-like tissue with an abundance of COL fibers surrounding the PLLA pores. The scaffold implanted with a layer of HAp covering its pores was invaded by tissue that showed zones with an aspect very similar to osteoids (Fig. 4c, asterisks), suggesting regenerated tissue more mature than in the other scaffolds. Quantification of the area occupied by these osteoid structures is shown in Fig. 4d and the extent of the COL type I expression (see [111]) confirmed these observations.
Fig. 4

Masson’s trichrome staining of the histological sections of tissue regenerated by PLLA, PLLA/HAp, and PLLA/HAp/SBF scaffolds 6 weeks after implantation: (a) View of the PLLA implant inside the structure of the subchondral bone. (b) Detail of the edge between the implant and host bone tissue. (c) Detail of the scaffolds and tissue regenerated inside their pores. (D) Percentage area occupied by osteoids inside the scaffolds, as determined by ImageJ software (10 fields per sample were used for the quantification). For the analysis, p < 0.05 was considered as significant. Implants were placed in lesions performed in the femoral condyle of 3-month-old healthy sheep. Tissue filling the pores of the scaffolds can be observed, but osteoids (asterisk) are only present in the PLLA/HAp/SBF scaffold. Scale bars: 200 μm (a), 100 μm (b, c). Reprinted with permission from [111]

Despite the fact that organic–inorganic composites of PLA are beneficial for bone regeneration in terms of bioactivity, the use of hydrophilic inorganic fillers such as HAp or SiO2 has certain drawbacks because of poor interaction with the hydrophobic PLA matrix. Even if these fillers are well dispersed within the matrix, the mechanical properties of the composite scaffolds tend to be equal or inferior to those of pure PLA. Filler particle surface modification by grafting prior to combination with PLA has been proposed as a way to overcome these drawbacks [112, 113]. Chemical grafting is tedious and expensive, but there are cheaper strategies that are also easier to apply and convenient. A mixture of the hydrophilic fillers HAp and SiO2 was more effective in enhancing the mechanical properties of PLLA scaffolds than either filler alone [114]. This ability was a result of the hydrophilic interaction between particles that were oriented such that their hydrophobic part faced the PLLA matrix, causing a stronger hydrophobic interaction with the polymeric matrix [114]. Consequently, the Young’s modulus of the scaffold containing a mixture of HAp and SiO2 was 30% greater than for the pure PLLA and reached a value of ~7 MPa (Fig. 5), which is considerably higher than the usual values obtained in polymeric scaffolds based on PLA. As expected, compression moduli decreased after some time in SBF as a result of polymer degradation.
Fig. 5

Apparent Young’s moduli at compression of PLLA, PLLA/HAp 90:10, and PLLA/HAp/SiO2 90:9.5:0.5 scaffolds after various times of immersion in simulated body fluid. Reprinted with permission from [114]

3.5 Commercially Available Products for Skeletal Tissue Engineering

Despite the abundance of literature related to the use of PLA and its combinations in tissue engineering, very few products reach the market or clinical trials. In tissue engineering of cartilage, most of the matrices are composed of COL or hyaluronic acid (HA) [58]. BioSeed®-C (BioTissue AG, Zürich, Switzerland) is the only matrix material for autologous chondrocyte implantation based on PLA that is on the market. It is a PGA/PLA/poly-p-dioxanone supportive matrix (derived from one of the materials of biodegradable sutures, polyglactin 910) that is combined with culture-expanded autologous chondrocytes suspended in fibrin glue and used for the regeneration of cartilage lesions [115, 116]. The radiological outcome of the clinical implantation of BioSeed®-C in knee lesions seems to be better than the traditional therapy of autologous chondrocyte implantation, which is a proof of the benefits of using scaffolds in cartilage defects [115].

Alvelac is a porous, osteoconductive, biocompatible, and biodegradable synthetic scaffold synthesized from PLGA and polyvinyl alcohol (PVA) produced via 3D printing by Bio Scaffold International Pte Ltd. Its proposed application is for regeneration of alveolar bone after teeth extraction [117], although other orthopaedic uses are being explored. As it is produced by 3D printing, the advantage of this product is that scaffolds with the shape of the defect in the patient can be produced. Soft Tissue Regeneration Inc. is currently involved in a clinical trial study of their product L-C Ligament®, based on a PLLA scaffold designed to facilitate re-growth of a patient’s anterior cruciate ligament (ACL) within the knee [118].

4 Surface Modification of Polylactides by Chemical and Physical Methods

4.1 Surface Modification of Polylactides

Certain disadvantages of polylactide polymers in several tissue engineering applications have been briefly discussed, such as the lack of cell adhesion motives and their relative hydrophobicity, which can lead to undesired effects on protein adsorption and subsequent interaction with cells [8, 9]. Surface modification of polymers can be achieved by physical or chemical methods and has been reviewed in detail in a number of publications [40, 119]. Physical modification of PLA by adsorption of ECM proteins such as COL and FN has been suggested by some authors, although the process is not well defined and lacks strategies for sterilization because of the negative effects on proteins, making this method less suitable [71]. Other methods of adsorption such as the layer-by-layer (LbL) technique can be used to apply more stable molecules such as glycosaminoglycans (GAG) that can mimic the natural environment of cells, but can be sterilized by UV radiation or other techniques [120, 121]. A wide variety of chemical treatments for polymers have been described in more detail in other reviews (e.g., [122]). Treatment of polymers for the introduction of functional groups such as amino groups has been carried out using nitrogen plasma or allylamine [123, 124]. Plasma treatment is useful for polylactides, and its copolymers and blends, because the treatment does not involve any organic solvents that can migrate into the polymer or water that can induce hydrolysis. On the other hand, previous work has successfully shown the use of wet chemistry for modification of PLA with bioactive molecules such as heparin (HEP) or hirudin for applications in contact with blood [125, 126]. Extensive reviews on the surface modifications of PLA can be found elsewhere [71, 127]. Briefly, we describe ways to modify PLA by simple adsorption techniques or chemical modification to introduce amino groups that promote cell adhesion [53] and adsorption of bioactive GAG, which promotes not only cell adhesion and spreading, but also differentiation of hMSCs.

4.2 Covalent Activation versus Adsorptive Binding of Polyamines

Poly(ethylene imine) (PEI) is widely used as an agent to support transfection of cells and has been immobilized on polymers such as polyimides, showing promoting effects on the adhesion of keratinocytes [128]. Hence, surface modification of PLLA cast films with different PEIs of either high (750 kDa) or low (25 kDa) MW (either by adsorptive or covalent immobilization) was carried out to study the effect on the activity of bone cells. The samples obtained after adsorption of PEI were designated as AH (PEI MW 750 kDa) or AL (PEI MW 25 kDa), where A stands for adsorption and H and L for high and low MW PEI, respectively. Covalent binding of PEI to terminal carboxylic groups of PLLA was achieved using N-(3-dimethylaminopropyl)-N′-ethylcarbodiimide hydrochloride (EDC) and N-hydroxysuccinimide (NHS). Details of the procedure can be found elsewhere [129]. The samples were designated as CH (PEI MW 750 kDa) or CL (PEI MW 25 kDa), where C stands for chemical binding of PEI.

Because the surface composition and properties of biomaterials dictate the adsorption of proteins and adhesion of cells [16], surface characterization of substrata using X-ray photoelectron spectroscopy (XPS) was carried out to quantify the chemical composition after immobilization of PEI (as shown in Table 1). PLLA contained 61.6% carbon and 38.4% oxygen. After PEI modification, the ratio of C to O increased because carbon from PEI was bound to the surface. In addition, Table 1 shows that nitrogen was found after deposition of PEI and increased for samples modified with high MW (HMW) PEI, especially for sample CH with 3.7 mol% nitrogen, indicating that the covalent immobilization was most efficient. The amino group density on PEI-modified PLLA films was quantified by acid orange assay [128] and showed consistency with XPS studies regarding the higher amount of amino groups when HMW PEI was used (Table 1). Water contact angle (WCA) measurements using the sessile drop method confirmed these findings, showing a decrease in WCA from about 72° for relatively hydrophobic PLLA to about 57° for HMW PEI, irrespective of the modification method. The results were in line with previous studies on PEI immobilization using other types of polymers [128].
Table 1

Surface properties of PLLA films after immobilization of poly(ethylene imine)


Surface elemental composition (mol%)

Amino group density (×1016 cm−2)

Water contact angle (°)







72.4 ± 0.3





2.51 ± 0.19

60.5 ± 0.9





9.36 ± 0.62

59.4 ± 0.6





12.14 ± 0.40

56.8 ± 2.7





12.79 ± 0.36

56.2 ± 1.9

AL adsorptive binding of low molecular weight PEI), CL covalent binding of low molecular weight PEI, AH adsorptive binding of high molecular weight PEI, CH covalent binding of high molecular weight PEI.

Surface elemental composition was obtained by XPS, amino group density of PLLA surfaces measured by acid orange assay, and water contact angle with the sessile drop method (Modified from [129])

The promoting effect on biocompatibility and osteogenic activity of PLLA modification with PEI of different MWs by either adsorption or covalent binding was further studied with the osteoblast-like cell line MG-63 [130]. This cell line has a controlled growth behavior and expresses osteogenic features such as alkaline phosphatase (ALP) activity [131]. Studies of adhesion and proliferation of MG-63 cells on different sample surfaces revealed that adhesion and cell growth were similar for all surface modifications with PEI but showed significantly lower cell growth on plain PLLA, which corresponds to cell behavior on more hydrophobic materials [53].

The osteogenic activity of cells was quantified by the expression of ALP, an enzyme involved in the calcification of bone [131]. As shown in Fig. 6, no ALP activity of MG-63 cells was detected on the tested surfaces at day 1, which reflects the fact that ALP activity is upregulated at later stages of culture [132]. After 6 days, the ALP activity of MG-63 cells increased significantly and continued to rise until day 9. Cells cultured on PEI-modified samples expressed significantly higher ALP activity (p < 0.05) than those grown on plain PLLA. However, no remarkable differences were observed between the different PEI-modified surfaces.
Fig. 6

Alkaline phosphatase (ALP) assay for differentiation of MG-63 cells measured as absorbance (A405 nm) versus culture time. PLLA plain PLLA; AL or AH PLLA modified with low (L) or high (H) molecular weight PEI by adsorption (A); CL or CH PLLA modified with low (L) or high (H) molecular weight PEI by covalent binding (C). Asterisks indicate a significant difference between PEI-treated samples and PLLA, with p ≤ 0.05. Reprinted with permission from [129]

The results indicate that PEI is useful for covalent and noncovalent modification of polymer surfaces such as PLLA films. The inherent cytotoxicity of HMW PEI found in other studies [133, 134] was not observed here, which suggests that it is a suitable candidate for making hydrophobic polymer surfaces more biocompatible. The choice of immobilization method (adsorptive or covalent) and the MW of PEI (low or high) depends on the desired application, such as the preparation of polyelectrolyte multilayers (PEM) on implant materials.

4.3 Physical Surface Modification of PLLA with Polyelectrolyte Multilayers

Of the possible physical surface modification techniques, LbL assembly of polyelectrolytes (PEL) is a simple method for generating multilayers on charged material surfaces [120, 135]. The LbL technique is based on electrostatic interaction between charged substrata and PEL in aqueous solution, followed by ion pairing during their complexation on the surface [135]. In addition, hydrogen bonding and hydrophobic, host–guest, and other interactions are useful for complexation of macromolecules, particulate material, and cells on surfaces in an LbL approach [120]. The inner structure and surface properties of PEM strongly depend on PEL characteristics such as MW and presence of basic or acidic functionalities (strong or weak). Solution temperature and composition in terms of PEL concentration, ionic strength, and pH are also important [136, 137]. The LbL technique can also be applied to prepare multilayers that partly resemble the composition of the ECM by using matrix proteins and GAG. Here, we present an example using CHI as polycation and GEL, HA, and HEP as polyanions for formation of multilayer films on PLLA films. The experimental details can be found elsewhere [14]. The PLLA films were pre-activated with HMW PEI by covalent binding to maximize the content of amino groups that acquire a positive surface charge at acidic pH values used during PEL adsorption [129]. Multilayer formation was carried out until 10 or 11 single layers were obtained. Layer growth was studied by surface plasmon resonance (SPR), which quantified the adsorption of PEL by changes in the angle shift. Figure 7 shows the results of SPR measurements for GEL/CHI, HA/CHI, and HEP/CHI multilayer formation and represents the SPR angle shifts during formation of PEM on the sensor surface with a total of 11 PEL layers.
Fig. 7

Changes in SPR angle as a function of number of layers of gelatine/chitosan (GEL/CHI, filled circles), hyaluronan/chitosan (HA/CHI, open circles) and heparin/chitosan (HEP/CHI, filled squares) formed from 2 mg mL−1 PEL solutions in 0.14 M NaCl. Layer 1 represents PEI as primary layer; even layers polyanions (GEL, HA, HEP); odd layers CHI. Reprinted with permission from [14]

The SPR curves revealed different increases in layer mass for the three PEL pairs studied. Combination of the strong PEL HEP with the weak PEL CHI resulted in the largest angle shifts, indicating highest layer mass. Multilayer growth was exponential compared with the rather linear growth of HA/CHI multilayers. Furthermore, growth of the HA/CHI system was low, and almost insignificant for the GEL/CHI system. In addition, WCA measurements indicated whether the differences in growth regimes resulted in a specific surface wettability that depended on the PEL pair as well as the terminal layer. As seen in Fig. 8, plain PLLA films were rather hydrophobic, with a WCA value of 72°. Modification with PEI led to a decrease in WCA to 60°, which was in accordance with previous investigations [24, 129]. When this surface was exposed to the polyanions, a slight increase in WCA was observed, especially for the PEL pair HA/CHI (71°) and GEL/CHI (65°). The terminal layer of CHI/HEP had a lower WCA of about 50°. The results of these studies showed that both chemical and physical surface modification of PLLA can change the overall physical properties such as wettability, but also the chemical composition of the surface without a change in bulk properties of the material.
Fig. 8

Water contact angles (WCA) measured after each adsorption step during multilayer formation for PEL pairs of gelatine/chitosan (GEL/CHI, filled circles), hyaluronan/chitosan (HA/CHI, open circles) and heparin/chitosan (HEP/CHI, filled squares). Reprinted with permission from [14]

Human MSCs, isolated from bone marrow and other tissues, possess multipotent differentiation capacities that are stimulated by appropriate signals [138, 139]. The induction of MSC differentiation is a highly programmed, lineage-specific process at the molecular level controlled by hormones, cytokines, and growth factors [140] and by the nature and topography of scaffolds [141]. We studied whether the modification of PLLA with PEM, which leads to a change in the microenvironment because of the different surface composition, can have a synergistic effect in the presence of osteogenic stimuli on the osteogenic differentiation of MSCs. Because long-term culture in normal medium revealed strong differences in cell morphology and growth of large nodules on HA and HEP layers, alterations in osteogenic differentiation were expected [129, 142]. Here, osteogenic differentiation of hMSCs was visualized by staining calcified areas with Alizarin Red, as shown in Fig. 9.
Fig. 9

Alizarin Red staining of human mesenchymal stem cells (hMSCs) cultured on multilayers from gelatine/chitosan (GEL), hyaluronan/chitosan (HA) and heparin/chitosan (HEP) in normal medium (DMEM with 10% FBS, left) or osteogenic medium (DMEM, 10% FBS and osteogenic inducer, right) after 3 weeks. Note that cells were cultured on the terminal polyanion layers (Scale bar = 100 μm). Reprinted with permission from [14]

The strongest staining in cells cultured in normal medium was found in cultures on CHI/GEL multilayers (Fig. 9, left column), where large areas with strong red color were visible, indicating the formation of calcified areas. Previous reports have indicated that strong spreading of hMSCs resulted in upregulation of genes related to osteogenic differentiation [143]. The results in this study are consistent with previous reports [143], because both cell adhesion and proliferation studies have demonstrated strong spreading of hMSCs on CHI/GEL multilayers [144]. In contrast to this, hMSCs cultured on CHI/HEP grew in aggregates, but did not show significant staining. Weak staining was also observed in cultures on plain PLLA and CHI/HA, indicating limited formation of calcified areas. Interestingly, some weak staining was only found there in nodules formed during long-term culture of hMSCs.

Striking differences were found when hMSCs were cultured on PEM-coated PLLA in osteogenic medium (Fig. 9, right column). Generally, it is anticipated that nodule formation is a consequence of osteogenic differentiation, which seemed to be supported by the more hydrophilic nature of the polyanions HA and HEP. These studies on MSC differentiation demonstrate that the simple coating of PLLA with PEM made of ECM components can lead to a change in the microenvironment of cells, thus promoting the desired differentiation of MSCs into osteoblasts.

5 Summary and Outlook

We summarize here the benefits of using PLA in engineering of skeletal tissues. Bimodal pore architectures can be obtained from freeze-extraction and particle leaching with micro- and macroporosity that enable modulation of permeability and mechanical properties by changing the solvent content and porogen amount. Furthermore, several bulk modifications of PLA (copolymerization, blending, network formation, and composite design) allow the properties of scaffolds to be adapted for different requirements in engineering of bone and cartilage. Although the feasibility of PLLA scaffolds for the culture and differentiation of MSCs into chondrocyte-like cells has been demonstrated [57], combination of PLA with hydrogels seems to be a more powerful strategy for articular cartilage regeneration [103]. Organic–inorganic composite scaffolds of PLA can be optimized as structures for bone replacement, where the filler acts as mechanical reinforcement, confers bioactivity, and neutralizes acidic degradation products of PLA.

In addition to changing the bulk properties of PLA, different surface modification techniques can be applied to tailor the surfaces for specific applications in tissue engineering. We present simple physical (adsorptive) and chemical modifications that could allow improved adhesion and growth of bone and other cells. In addition, ECM components such as proteins and GAG can be used to change the microenvironment of cells, fostering their differentiation into the desired lineage, such as described here for osteogenic differentiation of hMSCs.



This work was supported by Marie Curie Industry-Academia Partnerships and Pathways (FP7-PEOPLE-2012-IAPP, with grant agreement PIAP-GA-2012-324386) and IOF-Marie Curie fellowship program (Protdel 331655) as well as the German Research Society (DFG) through Grant GR 1290/10-1 and the Spanish Ministry of Economy and Competitiveness through the MAT2016-76039-C4-1-R Project (including Feder funds).


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Copyright information

© Springer International Publishing AG 2017

Authors and Affiliations

  • Gloria Gallego Ferrer
    • 1
    • 2
  • Andrea Liedmann
    • 3
  • Marcus S. Niepel
    • 3
    • 4
  • Zhen-Mei Liu
    • 5
  • Thomas Groth
    • 3
    • 4
  1. 1.Centre for Biomaterials and Tissue Engineering (CBIT)Universitat Politècnica de ValènciaValenciaSpain
  2. 2.Biomedical Research Networking Center in Bioengineering, Biomaterials and Nanomedicine (CIBER-BBN)ZaragozaSpain
  3. 3.Biomedical Materials Group, Institute of PharmacyMartin Luther University Halle-WittenbergHalle (Saale)Germany
  4. 4.Interdisciplinary Center of Materials ScienceMartin Luther University Halle-WittenbergHalle (Saale)Germany
  5. 5.Faculty of DentistryUniversity of TorontoTorontoCanada

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