The use of engineered polymer scaffolds as structural bone substitutes offers a promising alternative to biological allografts. Scaffolds can be engineered to tune their chemistry and microarchitecture to optimize delivery of cells and osteogenic factors as well as to enhance their mechanical properties. The success of a scaffold as a structural bone substitute depends on achieving and maintaining a balance between the biological and mechanical (functional) requirements, both of which depend upon scaffold’s porosity.
In this study, we evaluated PLA scaffolds that had a volumetric porosity in the range of 30 – 40%, which was reduced by half with the inclusion of βTCP mineral particles in the scaffolds. These porosity values are significantly lower than the typical values reported in the literature. For example, Charles-Harris et al. (2007) analyzed the microarchitecture and mechanical properties of PLA and calcium phosphate glass scaffolds11
and reported porosity values in the range of 93–96.5%, and compressive moduli in the range of 0.06–0.20 MPa. Zhang et al. (1999) evaluated PLLA and PLGA scaffolds combined with hydroxyapatite particles and reported porosity values ranging from 85–95%, which compressive modulus and yield strength values in the range of 6–11 MPa, and 0.25-0.4 MPa, respectively.12
Lin et al. (2003) designed and characterized axially oriented PLDL scaffolds with porosity in the range of 58–80%, and average compressive modulus and ultimate strength values ranging from 43.5–168.3 MPa and 2.7–11.0 MPa, respectively.7
When compared to typical in vivo
ambulatory stress experienced by a mouse’s bone, which was estimated to be 16.8 MPa by multiplying the modulus of elasticity of a 12 week mouse femur 8
by the typical strains measured in vivo
it is likely that scaffolds with high porosity will fall short of the functional requirements in vivo
in this mouse model. In contrast, our PLA and PLA/βTCP scaffolds had average maximum compressive strength of 17.8 MPa and 35.7 MPa, which exceeds the estimated in vivo
stresses in a load bearing defect model in the mouse.
In addition to their functional load-bearing role, scaffolds can mediate delivery of osteogenic factors in bone tissue engineering. Osteogenic factors, such as BMP-2, have been proven to induce endochondral bone formation in vivo
For BMP-2 to be effective it must be delivered at an initial high dose to recruit osteoprogenitor cells and have a sustained release over an extended period to differentiate the progenitor cells. The retention dose and rate of release are determined by the microarchitecture and material characteristics of the scaffold as well as the factor entrapment technique. One common approach, which we followed in this study, is to load the protein by soaking the polymer scaffold in a buffered solution of the recombinant factor. The microarchitecture including porosity and pore size determine the surface area available for growth factor binding, while the interconnectivity of the pores can mediate uniform protein binding throughout the scaffold and can also permit fluid flow through the scaffold to enhance protein release by different mechanisms. The material characteristics of the scaffold can influence the affinity of the growth factor to the scaffold surface. Surface charge affects the binding and conformation of the bound proteins.15
In the case of PLGA polymers, the binding affinity of the factor is influenced by several material characteristics including the polymer acidity, the molecular weight, and the lactide/glycolide ratio.16
In this study, we used the soaking method to load our low porosity PLA and PLA/βTCP scaffolds with BMP-2 and VEGF and evaluated retention and release of these factors from the scaffolds in vitro
. Release kinetics from scaffolds can obviously be influenced by a number of factors, which in addition to the protein size (molecular weight or hydrodynamic radius) include polymer hydrophilicty (or hydrophobicity), degradation rate, swelling (water uptake) rate, and protein-polymer binding affinity. However, since our approach involved surface coating of the polymer scaffold rather than emulsion or dispersion of the protein in the polymer solution, the aqueous diffusion likely dominates release and is governed only by the molecular size of the protein17
as we expect no differential affinity or binding to the polymer. Since both BMP2 and VEGF form disulfide-linked homodimers, and their monomer molecular weights are equivalent (13 and 14 KDa, respectively), we therefore predicted that their aqueous diffusion coefficients are equivalent and assayed only for aqueous BMP-2 diffusive release. The retention (or entrapment) efficiency of the scaffolds was only 0.59 – 1.52% which may be related to the low surface porosity of the scaffolds. The fact that the porosity of the PLA scaffolds is nearly double the PLA/βTCP scaffolds is loosely correlated with the 2.6 fold increase in retention efficiency in the PLA scaffolds. To enhance the entrapment of rhBMP-2 on the scaffolds, we suspended the rhBMP-2 in type I collagen solution into which the scaffolds were subsequently soaked, to provide a thin collagen coat and enhance the retention efficiency. Type I collagen is the most abundant type of collagen and is present in bone. Alone, type I collagen has demonstrated little effect on bone formation but improved the bony interface between the graft and the host bone. Although, when used in conjunction with rhBMP-2, type I collagen has been shown to enhance bone formation in spinal fusions.18
Using this method we were able to increase the retention efficiency of the PLA and PLA/βTCP scaffolds by 2.1 fold and 5.3 fold, respectively, essentially equalizing the retention efficiency for both scaffolds. Since we did not attempt to gel the collagen, we posit that a thin layer of collagen coating likely was responsible for the modest enhancement of rhBMP-2 entrapment, and that effective infiltration of the collagen solution into the low porosity scaffolds was not achieved due to the viscous nature of the collagen solution. Alternatives to coating recombinant factors on the scaffolds should be considered including dissolving the osteogenic peptides in the polymers mixture during the scaffold fabrication. Such an approach has been shown to dramatically enhance the entrapment efficiency, reaching values in the range of 38 – 81% depending on the mixture formulation,19
and can potentially overcome the limitations of the soak and coat technique.
The majority (95–98%) of rhBMP-2 was released from the PLA and PLA/βTCP scaffolds in the first 12 hours, which is significantly faster than release kinetics reported by Burdick et al from 80% porous PLA scaffolds.20
The rapid release kinetics in our study supports the hypothesis that the rhBMP-2 retention was mainly due to surface coating rather than volumetric entrapment throughout the porous network. The kinetics of release of a growth factor is dictated by the mechanisms of release, which include diffusion-, swelling-, and erosion-controlled release.21
Diffusion-controlled systems mediate factor release from the non-degraded polymer due to a concentration gradient. Swelling-controlled systems enhance factor diffusion due to polymer swelling and opening of the intermolecular crosslinking, while erosion-controlled systems mediate factor release as a result of polymer surface erosion and degradation.21
Since PLA scaffolds undergo bulk erosion rather than surface erosion,3
the likely dominant mechanism of release in our scaffolds is diffusion. We empirically modeled the release kinetics using an exponential association relationship that conforms to the general form of asymptotic solutions of diffusion release described by Baker and Lonsdale22
for drug diffusion coefficient (D
) from a sphere of radius r
following the initial burst release phase (Mt
There was excellent agreement between the model and the experimental data (R2
range 0.87 – 0.99 for the individual samples) confirming that diffusion-controlled release was the dominant mechanism.
To evaluate the bioactivity of released growth factors in vivo
, PLA/βTCP scaffolds loaded with rhBMP-2 and rmVEGF120
were ectopically implanted in muscle pouches. That the scaffolds did not induce any detectable mineralization is likely due to the poor retention efficiency that effectively resulted in loading about 6 ng of either factor per scaffold. A survey of the literature suggests that a wide range of BMP-2 doses can be effective in inducing ectopic bone formation from as low as 100 to 500 milligram23
to as high as 50 microgram24
or even greater.25
However, since Aspenberg et al. demonstrated that a dose as low as 1 ng of rhBMP-2 was enough the elicit bone formation in porous coralline hydroxyapatite chambers implanted in rat tibiae,26
we elected to evaluate a lower concentration than typically used in these ectopic models. The low dose combined with the burst release may not have been sufficient to induce ectopic bone formation in our model. Unlike an orthotopic defect such as the tibia26
where osteoblasts and osteoprogentors are abundant, an ectopic (muscle) model might not afford enough osteoprogenitors to respond to rapid-release of low doses of rhBMP-2, which must still be verified by detailed analysis of the cellular and tissue histology in the milieu of the implanted scaffolds.
Contrary to the lack of osteogenesis in our ectopic model, a similar dose of rmVEGF120
resulted in a significant 1.8 fold increase in vessel volume in the milieu of the scaffolds. The vessels were mostly external to the scaffolds although some limited vascular ingrowth was observed within the scaffolds. Previous studies have shown that doses as low as 250 ng of rmVEGF120
induce ectopic vessel formation.27
The increased ectopic vessel formation with low doses of VEGF in our study and others is likely due to the availability of endothelial progenitors in the skeletal muscle. Angiogenesis is essential in the promotion of both endochondral and intramembraneous ossification but might not be on its own sufficient to enhance osteogenesis as it has been shown that combined delivery of BMP, VEGF and mesenchymal stem cells are required to significantly enhance ectopic bone formation relative to any single factor,28
while controlled, sequential VEGF and BMP-2 release from acellular composite polymers has been shown to enhance bone regeneration in ectopic implants and orthotopic defects.29