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The treatment of large bone defects resulting from trauma, inflammation or tumor resection still represents a significant and unsolved clinical problem in trauma and orthopedic surgery. The use of autologous tissue as gold standard is associated with a broad range of side effects and general drawbacks such as limited availability [1, 2].

The process of bone healing itself is highly complex and divided into four overlapping steps: (1) inflammation, (2) soft callus (3) hard callus and (4) bone remodeling. Each step is characterized by a specific set of cellular and molecular events [3]. So, at the molecular level, bone healing is driven by inflammatory cytokines, pro-osteogenic and angiogenic factors [4, 5]. However, the early stages of bone healing seem to be crucial for consolidation [6, 7].

Coating of biodegradable scaffolds or metallic implants has emerged as an important approach for improving biomaterial integration and bone regeneration. Because this integration is mediated by interactions between tissue cells and the implant surface, the prime goal is a biomaterial surface that can guide these interactions. The native composition of the organic extracellular matrix (ECM) of bone provides a rationale for this coating, since the ECM is not only a passive carrier but can also actively influence the cellular reaction [8] and biological activity of growth factors because of specific interactions with receptors and cytokines [9, 10]. It can therefore directly or indirectly influence migration and cell adhesion as well as proliferation and differentiation of bone cells [9–13].

In calcified bone the organic matrix consists of a highly ordered, site-specific network that is mainly composed of collagen (Coll) type I and smaller quantities of the glycoprotein (GP) fibronectin, the proteoglycans (PGs) decorin and biglycan as well as the glycosaminoglycans (GAGs) chondroitin sulfate (CS-A, about 90%), hyaluronan, dermatan and heparan sulfate [14, 15]. The additional presence of the non-collagenous components influences the structure and function of Coll fibrils in vivo.

Preparing an artificial ECM (aECM) as biomaterial coatings including these specific constituents therefore seems promising to mimic the natural microenvironment of bone cells in vivo and its ability to guide tissue repair. By including different ECM components it is feasible to tailor tissue-specific matrices with different structural and mechanical properties, bioadhesive character, susceptibility to enzymatic degradation and growth factor binding.

This review will focus on the generation of Coll-based, bone-related aECMs as coatings for scaffolds and implants. Furthermore the influence of these coatings on early and late bone formation is described.

Artificial ECM fabrication and characterization

Coll-based aECM can be prepared from suspensions of insoluble Coll fibers or solutions of Coll monomers, which form fibrils due to their spontaneous self-assembly termed in vitro-fibrillogenesis [16]. The majority is prepared from Coll type I since it is the most abundant Coll type in bone and easily obtained, for example, by acid extraction [17]. Performing in vitro-fibrillogenesis from Coll monomers including Coll-binding, non-collagenous components allows the modulation of aECM properties in terms of composition and structure in analogy to the in vivo situation.

The self-assembly of Coll fibrils takes place at physiological pH and temperature and is an entropy-driven process with the assembly pattern itself driven by charge distribution on the monomer surface [16]. Fibril formation and fibrillar structure are influenced by several factors including Coll type, Coll preparation (acid-extracted versus pepsin-digested), non-collagenous components, temperature, pH and buffer composition influencing the electrostatic interaction of monomers [18, 19]. In vitro-fibrillogenesis is in most cases performed according to the method of Williams [20] with conditions optimized to achieve a clear D-periodic banding pattern.

Multi-component aECMs can be generated by including different Coll types, GPs, PGs and GAGs in the fibrillogenesis process. Coll type I and Coll type III, for instance, form heterotypic fibrils [18, 21] and the glycoproteins fibronectin [18, 22] and laminin [23] can associate with Coll fibrils. PGs like decorin and biglycan can also be associated but to a limited degree since in general they strongly decrease the integration of Coll monomers and the fibril diameter [19, 24, 25].

Combinations of Coll and GAG are most commonly used for aECM preparation [26, 27]. Since GAGs, except for hyaluronan, are in general part of PGs, there is no direct analogy to the in vivo situation in this approach. However, their lower immunogenicity compared to PGs is an important advantage of their use in biomaterial coatings [28, 29]. In contrast to the non-collagenous ECM proteins the interaction of GAGs with Coll is rather unspecific and depending on the net negative charge of the polysaccharides. The ionic strength of the buffer system is important for formation of aECM from these components. At low ionic strength these Coll and GAGs combinations display fast aggregation and significantly affected fibril diameter [24, 30, 31].

Another option is to include different ECM components by adding them to pre-formed fibrils. In this case the interaction of non-collagenous-components was reported to depend on the Coll type as well as the experimental conditions under which the fibrils were formed [18].

After their formation Coll-based aECMs can be cross-linked to increase stability and reduce degradation rate. EDC [1-Ethylen-3-(3-dimethylaminopropyl) carbodiimide hydrochloride] is most commonly used [32]. Other options are glutaraldehyde, dehydrothermal crosslinking [33] and transglutaminase treatment [34]. The drawback of cross-linking however is a chemically altered material which might impair aECM functions. In addition there is further concern that residual chemical cross-linking agents might remain in the collagen matrix, leach out when implanted and be toxic to the host as demonstrated for glutaraldehyde [35].

Covalent immobilization of Coll to the biomaterial can be achieved via carbodiimide cross-linking with aminopropyl triethoxysilane and N,N´-disulphosuccinimidyl suberate as a linker [36], deposition on reactive maleic anhydride copolymers [37] or through low temperature glow discharge and cross-linking with glutaraldehyde [38]. The covalent approach has the advantage of leading to very stable coatings, but the biological components might be chemically altered since functional groups of proteins are part of the link.

Since fibrillar collagens exhibit a high surface to volume ratio leading to a high number of potential binding sites for the surface, adsorptive immobilization is a reasonable alternative for immobilizing these elongated macromolecules. The multitude of binding sites involved increases the total adsorption free energy and therefore desorption is minimal [39, 40]. This approach has the advantage of being very simple and can be conducted at physiological conditions. Drawbacks are that conformational changes cannot be excluded but are thought to be minimal.

Adsorption of Coll-based aECM is determined by the Coll share since its adsorbed amount was found to be independent from the non-collagenous ECM components included [31, 41], except for thin fibrils at high decorin concentrations [24].

Before multi-component Coll-based aECMs can be applied as biomaterial coatings they need to be biochemically characterized. Important aspects are their structure, their composition and the release of individual, non-covalently associated components as well as the functionality of these components. Another interesting aspect is the ability to interact with solute biological mediators (growth factors, cytokines), a feature attributed to many non-collagenous ECM macromolecules.

The amount of Coll type I integrated into fibrils, the quantity adsorbed to titanium and the amount of the GP fibronectin bound was found to be dependent on fibrillogenesis parameters such as ionic strength, concentration of phosphate and the presence of Coll type III [18]. Fibronectin release depended on the initially used fibronectin concentration. For concentrations below saturation about 20% desorbed over 5 days, while above saturation the excess was released within the first 2 h. For the immobilization of larger amounts cross-linking would therefore be necessary.

Including decorin in Coll type I matrices always resulted in a significant decrease of fibril diameter, while with CS-A/C there were only slight changes detectable [24] (Figure 1). The amount of decorin and CS-A/C bound and desorbed was found to be depending on their initial concentration and the ionic strength of the fibrillogenesis buffer.

Figure 1

Atomic force microscopy (AFM) images of bovine Coll type I fibrils after in vitro-fibrillogenesis in 30 mM phosphate buffer pH 7.4 of (A) pure Coll and (B) Coll with 5 µg chondroitin sulfate A/C (CS-A/C) per 100 µg Coll. Height (left) and amplitude (right) images are displayed; AFM tapping mode. Adapted from [42].

Furthermore the amount of GAG associated to Coll fibrils was found to be dependent on the Coll type, the Coll preparation and the GAG itself [19, 31, 41] suggesting an influence of the sugar backbone, sulfation degree and pattern. Of note, when GAGs are included in aECMs during in vitro fibrillogenesis of Coll without cross-linked there is always a significant desorption taking place in the first hour of incubation at physiological conditions [31, 41], while afterwards GAG release was only marginal. This initial release could be prevented by cross-linking GAG to preformed fibrils in which the immobilized amount is a function of free functional groups, the concentration of GAG and the cross-linking agent [43]. In a recent study it could be demonstrated that aECM coatings consisting of collagen and CS or sulfated hyaluronan are stably attached to dental implants since the collagen content was not diminished by the insertion process [44].

Interactions of Coll-based aECMs with growth factors were reported for pure Coll as well as for multi-component matrices. TGF-β, bFGF, HGF and PDGF-BB interact with pure Coll matrices, which stabilized them and retarded their release rate [45, 46]. BMP-2 was reported to interact much stronger with matrices from Coll and the PG perlecan than from those of pure Coll and also displayed a seven times slower release rate [47].

Matrices of Coll and heparin displayed an improved release kinetic of VEGF compared to bare titanium surfaces [48]. The initial burst within the first 24 h was diminished and from the third day on heparin-containing matrices showed a higher VEGF release than those of pure Coll matrices or bare titanium. Coll-based aECM containing CS-A/C or chemically sulfated hyaluronan likewise reduced the initial burst of TGF-β1 within the first 24 h in a sulfate dependent manner (high-sulfated > low-sulfated) compared to a pure Coll matrix [41]. Thereafter, only those matrices containing sulfated hyaluronan show a further TGF-β1 release. Coll synthesis of human mesenchymal stem cells (hMSC) was found to be promoted by pre-adsorbed TGF-β1 on all aECM in comparison to matrices without growth factor. The greatest effect was found with the aECM containing high-sulfated hyaluronan. However, there was no correlation to the amount of TGF- β1 adsorbed or released and the effect on collagen I synthesis, suggesting that the growth factor was only specifically recognized by the cells when adsorbed to matrices containing sulfated hyaluronan. This clearly demonstrates the potential of these coatings in influencing cellular behavior due to binding and presentation of biological mediators.

Moreover, chemically sulfated hyaluronan derivatives are interesting model substances for investigating the structure-function relationship of GAGs, since they can be synthesized with defined chain length and sulfation degree in contrast to the batch-to-batch variability of native sulfated GAGs. Additionally, in contrast to unmodified hyaluronan, they display sulfation dependent interactions with BMP-4 and OPG, biological mediator proteins relevant to bone healing, modulating their bioactivity [49, 50].

But even in the absence of added growth factors Coll-based aECM are able to direct the cellular behavior of bone related cells. Matrices of Coll type I are well known for promoting cellular adhesion and spreading as well as proliferation of ostoblasts [40, 51]. Coll matrices containing sulfated hyaluronan stimulated the osteogenic differentiation of hMSC by increasing alkaline phosphatase expression and activity as well as calcium phosphate deposition in comparison to those containing Coll with and without CS-A/C [52]. At the same time aECM containing sulfated hyaluronan and CS led to a significant inhibition of osteoclast differentiation and resorption, which was found to be depending on their degree of sulfation [53].

Whether these in vitro findings on Coll-based aECM result in a subsequent improvement of bone healing around biomaterials was further assessed in several animal studies for Coll and Coll/CS-A coatings in vivo, while there is limited information in this respect on Coll matrices containing sulfated hyaluronan [44, 54].

Before in vivo application, collagenous matrices need to be sterilized. Finding an effective sterilization procedure for Coll-based aECMs that will not destroy their structural integrity, however, is a major challenge. To date there are no techniques at hand that are without certain drawbacks. Simple autoclaving is not applicable due to terminal denaturation of collagen [55]. Besides being a cytotoxic and carcinogenic compound, sterilization with ethylenoxide is known to chemically alter collagen/GAG matrices due to its irreversible reaction with free amino groups [56, 57]. Gamma-irradiation decreases the mechanical properties of collagen/chondroitin sulfate materials [58] and increases the degradation of collagen, as does electron beam (beta-) irradiation [59], which is attributed to the cleavage of peptide bonds [55]. Despite its damaging effects gamma-irradiation seems to be the best sterilization choice for retaining the bone-inductive properties of collagen-based biomaterials. In a comparative study using five different sterilization methods for demineralized rat bone, gamma-irradiation reduced the bone inductive capacity of demineralized bone by only 50%, while ethylenoxide on the contrary destroyed almost all bone-inductive capacity [60].

In vivo-analysis of Coll type I coatings

Titanium and its alloys are attractive materials for orthopedic applications displaying a high biocompatibility which results in good osteointegration with high bone-implant contact [61]. Coating of Ti pins with Coll resulted in a more intense cellular reaction in the early stages of bone healing around these Coll-coated compared to uncoated pins in a rat tibia model [62, 63]. Osteopontin and osteonectin, both non-collagenous proteins of the ECM and involved in bone healing, appeared earlier around Coll-coated Ti pins indicating an earlier onset of the bone remodeling process compared to uncoated pins (Figure 2). Both uncoated and coated implants were surrounded by newly formed lamellar bone after 28 days. The direct bone-implant contact and the amount of newly formed bone around the pin were increased around Coll-coated compared to uncoated pins without reaching statistical significance [62]. The fact that Coll affects the early stages of bone healing was also confirmed by a study with Coll-coated Ti implants (external fixator pins) in a sheep tibia model. The apparent new bone formation and increased activity of osteoblasts around the external fixator pins suggested an increased bone remodeling around these implants after 6 weeks [64].

Figure 2

Collagen-coated (Ti/Coll) and uncoated Ti implants placed as intramedullary rods in the rat tibia [62]. (A) Average count of cells stained against osteopontin directly at the implant surface per low power field (significances *p<0.05). (B and C) Immunohistochemical staining for osteopontin (red, white arrows) at day 7 after implantation (original magnification ×100).

Other groups also confirmed the benefit of coating Ti implants with Coll. In osteoporotic rats the modification with Coll improved the osteogenicity of these implants [65]. The amount of newly formed bone in cavities of Coll-coated cylindric Ti implants was higher than in cavities of uncoated Ti implants after 5 weeks in a goat model [66]. In contrast, no improved osseointegration was observed with porous Ti cylinders coated with Coll gel implanted into the sheep tibia [67]. However, different animal models and/or coating techniques make it difficult to directly compare these studies.

Hydroxyapatite (HA) is used in clinical practice as a bioactive and biocompatible material with excellent osteoconductive properties [68, 69]. It is well known that the combination of HA with Coll enhances the bioactivity of the implants by providing a source of calcium and phosphate ions that can be used by osteogenic cells to produce bone [70]. In a rat tibia model histomorphometric analysis showed that the bone-implant contact around HA/Coll implants was significantly higher compared to pure HA implants [71]. Some authors suggest that the addition of Coll to HA implants can induce the ossification process in the early stages of healing [72–74]. This finding is supported by our own studies. Newly formed woven bone could be detected around HA/Coll but not around the pure HA implant at day 6 [71].

Biocompatible and resorbable polymers such as poly(ε-caprolactone) and poly(lactic acid) and their copolymers are often used for biomedical application in bone and cartilage repair [75], for drug delivery systems [76], and as surgical suture [77]. Polycaprolactone-co-lactide (tradename: PCL) fiber from Catgut company (Catgut GmbH, Germany) is a clinical used suture that was used to manufacture an embroidered 3-dimensional scaffold [78]. Functionalization with Coll resulted in an improved cell attachment and proliferation in vitro [79]. In a critical-size defect of the rat femur the amount of new matrix material was 1.5-fold higher in the defect zone in the PCL/Coll group compared to uncoated PCL after 12 weeks [78].

However, Coll also serves as a substratum for collagenases and matrix metalloproteinases [71, 80]. It therefore enhances bone resorption as well as bone formation resulting in an overall increased bone remodeling [62, 71].

In vivo-analysis of Coll type I/ chondroitin sulfate coatings

Besides Coll, CS seems to be a promising coating for biomaterials. It can enhance bone healing by mediating the binding of bone like cells as osteoblasts and capturing soluble molecules such as growth factors into the matrix [81]. CS was used in different animal studies. It could be demonstrated in a rat tibia model that Coll/CS-coating of Ti pins supports early healing. Whereas around uncoated pins just a primitive fibrin network with only a few macrophages was observed, a reparative granulation tissue with a high number of mononuclear macrophages was visible around Coll/CS-coated pins 4 days after implantation. Finally, the bone-implant contact around the coated Ti pins was significantly higher compared to pure Ti pins after 28 days. Compared to Coll coating alone, the early increase in osteoblasts and osteoclasts was even more pronounced with Coll/CS coating, indicating an additional effect through CS [63]. In a sheep experiment under loading conditions in the tibia the extraction torque of external fixation pins was slightly increased when pins were coated with Coll/CS in comparison to uncoated pins [64]. The favorable in vivo results of Coll/CS coatings on Ti pins were also confirmed by further studies examining dental implants [82–84]. In a minipig study the Coll/CS-coating showed significantly increased bone-implant contact compared to uncoated surfaces after 6 weeks of implantation [82]. In a further study this was verified for Coll/CS-coated Ti implants compared to controls after 1 month. However, after 2 months no differences could be detected between these groups indicating that Coll/CS enhances early bone formation [83]. Nevertheless, studies in osteoporotic rats could not show significant differences in peri-implant bone formation of Coll/CS-coating in comparison to uncoated Ti pins after 2 and 4 weeks. It is therefore assumed that the coatings may not be able to exert their beneficial effects around a presumably reduced osteoblastic precursor pool [84]. Therefore, long-term studies in animal models with reduced bone healing capacity are necessary to completely evaluate the potential of such coatings.

Findings on the effects of the addition of CS to Coll-coated HA implants further emphasize its beneficial effect on early stages of bone healing. Non-resorbable HA/Coll implants induced a significantly higher number of cathepsin D-positive and TRAP (Tartrate-resistant acid phosphatase)-positive cells around implants in the rat tibia after 4 and 7 days when coated with CS [85]. Cathepsin D is an important marker for cells of the monocyte/macrophage lineage and is known to play a role in bone remodeling in both physiological and pathological conditions. It was found in chondroclasts, chondrocytes, and osteoblasts [86]. These cells are required for a successful tissue turnover in the interface region, reflecting a higher bone remodeling activity around HA/Coll/CS implants compared to controls [85].

In a critical midshaft defect of the sheep tibia the callus appeared earlier around HA/Coll/CS implants than around HA/Coll implants. In addition, volume and amount of newly formed bone were increased around HA/Coll/CS compared to HA/Coll [87].

Coll/CS-coated PCL scaffolds led to a 9.5-fold increase in hMSC number over 21 days in vitro (Figure 3A). In a femoral critical-size defect in rat the effect of this Coll/CS-coating on PCL implants was investigated. In this group the highest new bone formation compared to uncoated implants was observed [78]. Further, the matrix deposition was 1.7-fold higher compared to the uncoated group (Figure 3B). Subsequently the coated PCL scaffolds were investigated in a pilot study in sheep. After 3 months new bone formation of 63% and after 12 months of 172% compared to the contralateral tibiae was observed indicating an osteogenic stimulation and induction of bone formation by the coating (Figure 3E) [88].

Figure 3

(A) Proliferation of human mesenchymal stem cells on pure PCL and Coll/CS-coated PCL scaffolds over 21 days, initially seeded 3000 cells per scaffold. (B) Quantification of new matrix deposition in femoral critical-size defect in rat after 12 weeks (significances *p<0.05). (C and D) Representative images of Masson-Goldner trichrome stained sections of the defect zone after 12 weeks; uncoated PCL (C), Coll/CS-coated PCL (D); ×2.5. (E) Quantification of the new bone formation after 3 and 12 months. The corresponding section of the contralateral tibiae were used as reference and set 100%. All data are represented as mean±SD.

The positive effects of Coll/CS-coatings on polymers were also confirmed by other authors. Vandrovcová et al. [89] investigated the influence of a Coll/CS-coating on poly(lactide-co-glycolide) (PLGA) scaffolds on the cellular behavior of osteoblast-like cells in vitro. After 3 and 7 days in culture the cell number was significantly increased in comparison to PLGA alone or PLGA coated with Coll [89].

In summary, all data indicate that CS could have additional effects compared to pure Coll-coatings. It has been shown in several in vitro studies that GAG-modified implants influence adhesion, proliferation and differentiation of mesenchymal stromal cells and osteoblasts [90–92]. Therefore, GAGs are actively participating in the bone healing process. Nakamura et al. [93] showed that GAG not only act as mechanical interlock but also have molecular functions. Degradation of GAG on Ti implants resulted in a decreased interfacial strength indicating the specific role of GAGs in the establishment of mineralized tissue-titanium interfacial adhesion [93]. Their sulfate groups as well as their carboxyl groups are suggested to interact with mineral structures in bone such as HA [94, 95]. CS is able to bind and modulate proteins like tumor necrosis factor or fibronectin [96–98] which might result in an enhanced attachment of cells on the implant surface [95] and an improved healing process.


Coating of biomaterials with aECMs offers a great potential to construct a defined, tissue-inducing microenvironment which in turn will improve bone healing and implant stability. Indeed several studies demonstrated that Coll/CS-coating enhanced new bone formation in small and large animal models. Nevertheless, an increasing knowledge about the role of ECM components for proliferation and differentiation of cells as well as tissue regeneration will lead to defined adjustments of aECM coatings to the required properties for the healing process. A promising approach therefore is the use of GAGs with modified sulfation pattern, since it was shown that the binding affinities of bone-related growth factors depend on it [49]. This could result in a prolonged availability of the growth factor and hence an improved healing.


The authors would like to acknowledge the German Research Council (DFG SFB-TR67) and the CRTD.


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