Bone regeneration from human mesenchymal stem cells on porous hydroxyapatite-PLGA-collagen bioactive polymer scaffolds

Abstract

We have developed a novel multicomponent nano-hydroxyapatite-poly(D,L-lactide-co-glycolide)-collagen biomaterial (nHAP-PLGA-collagen) with mechanical properties similar to human cancellous bone. To demonstrate the bone forming capacity of nHAP-PLGA-collagen prior to in vivo experiments, nHAP-PLGA-collagen films and 3D porous scaffolds were seeded with human mesenchymal stem cells (hMSCs) to characterize cell proliferation and osteogenic differentiation. Over 21 days hMSCs seeded on 2D nHAP-PLGA-collagen films proliferate, form nodules, deposit mineral and express high alkaline phosphatase activity (ALP) indicating commitment of hMSCs towards osteogenic lineage. When seeded in 3D scaffolds, hMSCs migrate throughout the connected porous network of the nHAP-PLGA-collagen scaffold and proliferate to fill the scaffold voids. Over 35 days, cells express ALP, osteocalcin and deposit minerals with kinetics similar to osteogenesis in vivo. Adipogenic or chondrogenic differentiation is not detected in 3D constructs, indicating that in an osteogenic environment the presence of bone ECM specific molecules in nHAP-PLGA-collagen scaffolds support homogeneous bone tissue development. This ability of nHAP-PLGA-collagen matrices to provide biochemical stimulation to support osteogenesis from stem cells along with its high mechanical strength suggests that nHAP-PLGA-collagen is a suitable biomaterial for bone regeneration. This platform technology of covalently attaching ECM proteins and molecules with synthetic and natural polymers to adjust material properties and biochemical signaling has a potential for a wider range of applications in tissue engineering and regenerative medicine.

Keywords

Bone tissue engineering regenerative medicine collagen PLGA hydroxyapatite osteogenic differentiation

1. Introduction

A biomaterial that delivers biochemical signals to promote bone regeneration from stem cells while providing mechanical support during bone healing remains elusive. An effective graft biomaterial for bone regeneration must support load bearing during healing [1], recruit osteoprogenitor cells and promote osteogenic cell differentiation leading to new bone formation [2–4]. Ideally, the biomaterial resorption kinetics match the new bone formation rate to provide continuous mechanical support during tissue regeneration and healing. The complexity of natural bone requires development of intricate multicomponent bone extracellular matrix (ECM) substitutes [1,5] that provide structural support and biological signals necessary to promote bone regeneration during healing [6,7]. We recently developed a multicomponent covalently-linked biodegradable biomaterial, called nHAP-PLGA-collagen [8], with mechanical properties similar to cancellous bone that maintains high mechanical strength in an aqueous environment [9]. nHAP-PLGA-collagen consists of PLGA polymer which provides a strong biodegradable backbone; osteoinductive and osteoconductive hydroxyapatite bioceramic [10–13]; and collagen to provide biological stimulation for cell proliferation, osteogenic differentiation and bone formation similar to bone ECM in vivo [14,15]. Based on its favorable mechanical properties a preliminary assessment of the efficacy of nHAP-PLGA-collagen to provide necessary biochemical stimulation for stem cell proliferation and osteogenic differentiation in vitro is appropriate prior to validating the biomaterial in pre-clinical in vivo studies. In the present study, we quantify stem cell proliferation and osteogenic differentiation on two-dimensional films of nHAP-PLGA-collagen. Based on promising results using 2D films, three-dimensional porous nHAP-PLGA-collagen scaffolds were fabricated to characterize the kinetics of hMSC attachment, migration and osteogenic differentiation and demonstrate the usefulness of porous nHAP-PLGA-collagen scaffolds for bone formation from hMSCs.

2. Materials and methods

2.1. Preparation of two-dimensional films for cell culture

Synthesis of PLGA, nHAP-PLGA and nHAP-PLGA-collagen and processing them to produce 2D films are described in our previous studies [8,9]. Briefly, PLGA was synthesized by ring opening polymerization of 75 mol% D,L-lactide and 25 mol% glycolide monomers. nHAP-PLGA and nHAP-PLGA-collagen were synthesized using nanoparticles of hydroxyapatite as initiator. Collagen was crosslinked to nHAP-PLGA by modifying with N-hydroxysuccinimide (NHS) and N,N-dicyclohexylcarbodiimide (DCC) to obtain nHAP-PLGA-collagen. The relative mass ratio of nHAP:PLGA:collagen was 1.4:84.3:14.3. PLGA, nHAP-PLGA and nHAP-PLGA-collagen films with thickness of 300–400 μm were prepared by melt pressing [8]. Flat 2D disks of nHAP-PLGA-collagen were obtained from the films using 11 mm diameter surgical punch, which matches the diameter of the wells in a 48 well tissue culture plastic (TCP) plate (Fisher Scientific, Waltham, MA). Prior to cell seeding, 2D films were sterilized using 70% ethanol solution for 30 minutes and exposure to UV light for 4 hours [9]. Scaffolds were washed three times with Dulbecco’s phosphate buffered saline (DPBS without calcium and magnesium; Mediatech Inc., Manassas, VA) and kept in a fume hood overnight to evaporate the residual ethanol.

2.2. Cell culture and cell seeding

Human mesenchymal stem cells (hMSCs) isolated from bone marrow was purchased from Lonza Inc. (Walkersville, MD). hMSCs were cultured in T-75 cell culture flasks in 15 mL proliferation media consisting of Dulbecco’s modified Eagle’s medium (DMEM; Mediatech) with 10% fetal bovine serum (FBS; Atlanta Biologicals, Atlanta, GA), and 1% of a solution of amphotericin B, penicillin, and streptomycin (Mediatech) at 37°C in a 95% relative humidity 5% CO₂ cell culture incubator. Cells were fed fresh proliferation media every 3 days [16]. hMSCs were passaged using a 0.05% trypsin/EDTA solution into T75 flasks or seeded onto nHAP-PLGA-collagen films or scaffolds as described below. hMSCs within passage numbers 7–9 were used for all experiments.

For osteogenic differentiation studies, hMSCs seeded onto PLGA, nHAP-PLGA, nHAP-PLGA-collagen, collagen or TCP were cultured in osteogenic differentiation media [16,17]; proliferation media supplemented with 100 nM dexamethasone, 10 mM β-glycerol phosphate and 0.05 mM ascorbic acid (Sigma Aldrich, St. Louis, MO). An hMSC suspension of 100 μL in proliferation media at a density of $3.8 \times 10^{5}$ cells/mL was pipetted onto each 2D film (corresponding to 37,500 cells/cm²). Cells were allowed to attach for 8 hours in a 37°C 5% CO₂/95% air incubator and then the proliferation media was replaced with 300 μL of fresh differentiation media. Differentiation media was changed every 3 days.

2.3. Cell proliferation on 2D films

For cell proliferation studies, 100 μL hMSC suspension at $9 \times 10^{4}$ cells/mL were pipetted onto 2D films of PLGA, nHAP-PLGA or nHAP-PLGA-collagen in wells of a 48-well plate; corresponding to a seeding density of 9,375 cells/cm². This cell density was chosen only for the cell proliferation studies to ensure that cells did not reach confluence within 5 days of culture. Cell attachment efficiency was defined as the percentage of seeded cells that were attached on the film after 1 day. hMSCs seeded in TCP wells with and without collagen coating served as controls. TCP wells were coated with collagen at a concentration of 50 μg/cm² by pipetting 500 μL of a stock solution of calf skin type I collagen (Elastin Products Company, Owensville, MO) at 100 μg/mL in DPBS with 1 mM HCl into each well and incubating at 37°C for 4h [18–20] prior to cell seeding. Excess collagen was removed by triplicate washes with DPBS and air-dried in the cell culture hood for overnight.

Cells were treated with alamar blue (Molecular Probes Inc., Eugene, OR) after 1, 2, 4, 5, 6 and 7 days and incubated at 37°C for 4 h to allow conversion of resazurin to resofurin via reduction reaction of metabolically active cells. Absorbance values were obtained at 570 nm wavelength, using 600 nm as a reference wavelength using a microplate reader (Synergy HT, Biotek instruments, Winooski, VT). A calibration curve was prepared based on known numbers of cells to convert the absorbance values to cell number, which was used to extrapolate the number of cells for each condition at each time point.

2.4. Alkaline phosphatase activity on 2D film

After 7, 14, 21 and 28 days of 2D culture, cells were washed with DPBS and lysed with RIPA lysis buffer (Thermo scientific, Waltham, MA). The total DNA content of cell lysates was measured at each time point using the PicoGreen assay (Molecular probes). Alkaline phosphatase (ALP) activity of the cells was measured using an ALP assay kit (Abnova, Walnut, CA). ALP level was normalized to the total amount of DNA measured by PicoGreen assay at each time point and reported as amount of 4-Methylumbelliferone (4-MU) produced per μg of DNA/hour of reaction.

2.5. Evaluation of mineral deposition on 2D film using von Kossa stain

After 7, 14 and 28 weeks of culture, cells were fixed in 10% neutral buffered formalin solution (Fisher), and stained with von Kossa stain (American MasterTech, Lodi, CA) to qualitatively assess the mineral deposition. Von Kossa staining uses silver ions to react with phosphate groups in the matrix and produces brownish black precipitate, which serves as an indicator of mineral deposition [21,22]. Cells were counter stained with nuclear fast red for visualization. Since the biomaterial contains hydroxyapatite, a blank scaffold (without cells) was used as control and stained with von Kossa and nuclear fast red stain after incubation in cell culture media under the same conditions for 7 days.

2.6. Preparation of three dimensional scaffolds, porosity measurement and scaffold seeding

To obtain 3D scaffolds with 80–90% porosity, 10 wt% of melted nHAP-PLGA-collagen was mixed thoroughly with 90 wt% sodium chloride with a particle size of +80 mesh (177–210 μm) (Sigma Aldrich). The biomaterial/salt mixture was melt pressed to produce nHAP-PLGA-collagen/sodium chloride composite scaffolds with thickness of 700–800 μm. Scaffolds were soaked in distilled water for 24 hours with water change every 6 hours to leach out the salt [23] and vacuum dried for 24h to remove residual water [24]. Porous scaffolds were cut into 11 mm diameter discs using a surgical punch and stored in a desiccator under vacuum until use. Before using for cell culture, scaffolds were UV sterilized for 4 hours.

Photographs of the top surface of 3D porous scaffolds were taken using a Nikon SMZ800 stereo microscope (Nikon Instruments, Melville, NY). Three regions with a dimension of 2 mm × 1.5 mm each containing at least 25 pores were visualized to calculate the pore size distribution and average pore density using ImageJ software. Pore density was calculated as the ratio of total pore surface area to the total area of the region.

hMSCs were suspended in proliferation media at a concentration of 10⁷ cells/mL. One hundred microliters of cell suspension were pipetted on top of each scaffold in the well of a 48 well tissue culture plate and incubated for 8 hours at 37°C in a 95% air/5% CO₂ incubator to allow cells to attach to the 3D scaffolds. This corresponds to a seeding density of 10⁶ cells per scaffold. After 8 hours 400 μL of fresh differentiation media was added to each well. Differentiation media was exchanged every 3 days.

2.7. Histology of 3D tissue constructs

hMSCs cultured on 3D nHAP-PLGA-collagen scaffolds were fixed with 10% neutral buffered formalin (Fisher) after 7, 14, 21, 28 or 35 days of culture. Scaffolds were soaked in a 20% w/v sucrose solution in DPBS to allow infusion of embedding media to prevent crystallization [25], embedded in HistoPrep frozen tissue embedding media (Fisher), frozen in liquid nitrogen and stored at −80°C until cryosectioning. Tissue sections 40–70 μm thick were obtained by cryosectioning and attached to glass slides using a tissue capture pen adhesive (American MasterTech, CA). The 7 day time point samples were stained with hematoxylin and eosin (H&E) stain to visualize cell attachment and migration. Samples for 7, 14, 21, 28 and 35 days were stained with von Kossa to assess mineral deposition, and counterstained with nuclear fast red to visualize the cells. To assess homogeneity of hMSC differentiation, constructs cultured for 28 days were stained with alcian blue to identify the presence of acidic polysaccharides, indicative of chondrogenic differentiation [26,27] or with oil red O to check for adipogenesis [28,29].

2.8. Immunohistochemistry

3D porous nHAP-PLGA-collagen scaffolds seeded with 10⁶ hMSCs were fixed with 10% neutral buffered formalin solution (Fisher) after 7, 14, 21, 28 or 35 days of culture and embedded in paraffin wax using standard procedure [30]; except Histoclear (Electron Microscopy Sciences, Hatfield, PA) was substituted for xylene [31]. Sections of thickness of 7–10 μm were obtained using a microtome (Leica, Wetzlar, Germany), baked at 60°C overnight and deparaffinized using Histoclear. After rinsing with distilled water, sections were treated with hydrogen peroxide for 20 minutes followed by goat serum for 30 minutes to block endogenous hydrogen peroxidase and non-specific binding. Sections were incubated overnight with polyclonal antibodies against ALP or osteocalcin (Abcam, Cambridge, MA) followed by secondary antibody conjugated with horseradish peroxidase for 60 minutes. Images were taken using Olympus BX51 microscope (Olympus Corporation, Tokyo, Japan) and Bioquant imaging software.

2.9. Statistical analysis

Data are presented as mean ± standard error obtained from 3 independent experiments ( $N = 3$ ) with at least 3 samples for each condition ( $n = 3$ ). Cell proliferation and ALP activity values were compared for different materials at each time point separately using a one way analysis of variance (ANOVA) followed by pair-wise comparison with Tukey’s HSD test and Bonferroni test. All analyses were performed with IBM SPSS statistics software package. For all comparisons, the level of significance was $p ⩽ 0.05$ .

3. Results

3.1. Cell proliferation and osteogenic differentiation on 2D nHAP-PLGA-collagen films

One day after seeding, hMSC attachment efficiency was ⩾85% and cell density was between $8 \times 10^{3}$ to $12 \times 10^{3}$ cells/cm² for each biomaterial (Fig. 1(A)). Cell density increased more than 7-fold on nHAP-PLGA-collagen over 7 days (Fig. 1(A)) at which time cells reached confluence. ALP activity of hMSCs cultured on nHAP-PLGA-collagen films, a measure of hMSC commitment towards osteoblastic lineage [32], increases by 14 days and is maximum at 21 days (Fig. 1(B)). hMSCs begin to form colonies on nHAP-PLGA-collagen after 14 days (Fig. 2(C)), but mineral deposition was not prominent until 28 days of culture (Fig. 2(D)).

Fig. 1.

hMSC proliferation and ALP activity on 2D films. Data are (A) cell density and (B) ALP activity (expressed as 4-MU nmol/μg of DNA/hour) as a function of time for hMSCs cultured on tissue culture plastic, PLGA, nHAP-PLGA, nHAP-PLGA-collagen and collagen coated tissue culture plastic. Data are mean ± standard error of three independent experiments. ^∗Indicates significant difference ( $p < 0.05$ ) compared to TCP at the same time point.

Fig. 2.

Mineral deposition on 2D nHAP-PLGA-collagen films. Von Kossa stained cultures (A) after 7 days without cells, (B) after 7 days with cells, (C) after 14 days with cells, and (D) after 28 days with cells. There was no sign of mineral deposition after 7 or 14 days. Cells formed nodular colonies (circled in B) after 14 days. After 28 days (D), mineral deposition, was prominent (dark brownish precipitate outlined by dashed lines). Hydroxyapatite clusters are visible as dispersed dark spots (A) for von Kossa stained scaffolds without cells, which was used as negative control. Comparing the scaffold without cells (A) to the scaffold with cells after culturing for 28 days (D), it is evident that cells deposit mineral.

3.2. Cell proliferation and osteogenic differentiation on 3D nHAP-PLGA-collagen scaffolds

Porous nHAP-PLGA-collagen scaffolds have a uniform and interconnected three dimensional porous structures with average pore size of 270 ± 60 μm and average pore density of $81 \pm 2 %$ that extends throughout the construct thickness (Fig. 3(A–C)). At a seeding density of $1 \times 10^{6}$ cells/scaffold, cells proliferate and begin to fill the void spaces within 7 days of culture (Fig. 3(D)). Alkaline phosphatase is detected within 7 days of culture (Fig. 4(A)), increases over 14 days (Fig. 4(B)) and then declines for the rest of the culture period (Fig. 4(C–E)). Osteocalcin expression is also transient, with signal first detected at 21 days (Fig. 4(F–H)) that increases at 28 days (Fig. 4(I)) and 35 days (Fig. 4(J)).

Fig. 3.

Salt leached 3D scaffolds of nHAP-PLGA-collagen. (A) 3D view, (B) top surface, (C) cross section. Scaffolds have an interconnected three-dimensional porous structure with average pore size of about 270 ± 60 μm. (D) Cross section of the scaffold seeded with hMSCs and stained with H&E after 7 days of culture (arrow heads).

Fig. 4.

Kinetics of alkaline phosphatase and osteocalcin expression in nHAP-PLGA-collagen 3D scaffolds. Alkaline phosphatase (A–E) and osteocalcin (F–J) staining at the indicated time points. Red arrowheads indicate positive alkaline phosphatase after 7 days (A) and 14 days (B). Positive staining for osteocalcin is evident after 21 days (H), 28 days (I) and 35 days (J) (Red arrows). Scale bar represents 100 μm.

Fig. 5.

Mineral deposition in nHAP-PLGA-collagen 3D scaffolds. Cryosectioned scaffolds (A) without cells, after (B) 7 days, (C) 14 days, (D) 21 days, (E) 28 days and (F) 35 days were stained with von Kossa stain. Yellow arrows identify regions of brown precipitate representing mineral deposition. Scaffolds without cells serve as a negative control (A). Green arrows indicate the scaffold structure. After 35 days culture in osteogenic differentiation medium, scaffolds were negative for alcian blue (G) and Oil red O (H). Scale bar represents 200 μm.

Figure 5(B–F) shows the progression of mineral deposition in porous nHAP-PLGA-collagen scaffolds, confirming formation of a mature osteogenic phenotype from hMSCs over 35 days. Lack of proteoglycan (Fig. 5(G)) or lipid (Fig. 5(H)) expression after 35 days demonstrates that hMSC differentiation on 3D nHAP-PLGA-collagen scaffolds is homogeneous with little evidence of chondrogenesis or adipogenesis.

4. Discussion

Synthetic biomaterials for bone repair should provide mechanical support and biological functionality to promote bone tissue regeneration throughout healing. nHAP-PLGA-collagen, which incorporates the bone specific ECM components hydroxyapatite mineral and collagen polymer important for bone development and function [10,13,15], supports hMSC growth and osteogenic differentiation. hMSC proliferation on nHAP-PLGA-collagen films is robust and comparable to proliferation on collagen, a widely used substrate for hMSC attachment and proliferation [33,34]. Demonstration of osteogenesis within 3D scaffolds with kinetics similar to bone formation in vivo demonstrates the suitability of nHAP-PLGA-collagen as a bone regeneration biomaterial.

hMSC expression of ALP activity on nHAP-PLGA-collagen shows the commitment of hMSCs towards osteogenic differentiation [32,35], that is more robust than on pure collagen (Fig. 1(B)). The subsequent decline in ALP activity coupled with increased mineral deposition is consistent with hMSCs differentiation and maturation into osteocytes [36–38]. The 1.7 and 2.6-fold increase in ALP activity after 14 and 21 days of culture on nHAP-PLGA-collagen is higher than that observed in similar studies using different osteoconductive bone biomaterials such as calcium phosphate, hydroxyapatite alone or in combination with peptides [38–41], which indicates the combination of bone specific ECM protein, collagen type I and hydroxyapatite mineral had higher impact on initial commitment of MSCs towards osteogenic lineage.

Proliferation of hMSCs on nHAP-PLGA was slow (Fig. 1(A)), which might be attributed to the presence of bone specific mineral (nHAP) and with the absence of ECM protein (i.e. collagen), stem cells may have received signaling for early osteogenic commitment which might attenuate their proliferative capacity. Figure 1(B) also suggests that in the presence of nHAP in nHAP-PLGA (and also in nHAP-PLGA-collagen), osteogenic commitment of hMSCs were significantly high even though their initial proliferation was slow.

The uniform and interconnected three dimensional porous structure of nHAP-PLGA-collagen scaffolds provides a favorable environment for osteogenic differentiation of hMSCs [42,43]. ALP expression peaks and begins to decline earlier in 3D culture (Fig. 4) compared to 2D culture (Fig. 1B). Mineral deposition also begins earlier in 3D culture (Fig. 5) compared to 2D culture (Fig. 2). This suggests that matrix maturation and mineralization is faster in 3D culture, likely due to the close proximity and higher density of cells in 3D culture, which is essential for cell-cell signaling for phenotypic expression, tissue development and continued function [44].

In 3D culture, hMSCs migrate and fill scaffold voids within 7 days. ALP expression, which peaks at 14 days; osteocalcin expression, which is not detected until 21 days; and mineral deposition beginning at 28 days demonstrate that hMSCs progress through osteogenic differentiation and form mature bone cells. Minimal osteocalcin expression in 14 days (Fig. 4(F) and (G)) and transient ALP expression, is consistent with nodule formation [21,45] and matrix maturation [36,37] stages of osteogenesis. Expression of mature osteocyte phenotype similar to that observed during intramembranous osteogenic differentiation in vivo [46–48] confirms the ability of nHAP-PLGA-collagen to support bone formation from adult progenitor cells via intramembranous ossification.

hMSCs have multilineage differentiation potential and can differentiate into bone, cartilage, muscle, fat or other tissue [49,50]. Commitment of stem cells towards a specific tissue depends on scaffold compositions and mechanical properties as well as extracellular matrix proteins, nutrients, soluble and insoluble growth factors [51–53]. For clinical application of stem cells for tissue regeneration, homogeneous differentiation and expression of a single tissue lineage is critical [53]. hMSCs cultured in osteogenic media on 3D nHAP-PLGA-collagen scaffolds formed mature bone from progenitor cells (Figs 4 and 5) without expression of chondrogenic or adipogenic markers (Fig. 5). Incorporation of hydroxyapatite and collagen in the scaffold, two important biological factors for bone formation [10,13,15], likely enhance osteogenic differentiation of hMSC while preventing significant chondrogenic or adipogenic differentiation.

nHAP-PLGA-collagen has tensile strength comparable to human cancellous bone and significantly higher strength than pure collagen [54] or other bioactive osteogenic biomaterials like chitosan [55] and gelatin [56]. Our previous studies demonstrated that chemically bonded components of nHAP-PLGA-collagen provide superior mechanical properties in dry and wet condition and favorable interactions with cells while remodeling to maintain the strength and integrity of the scaffold [8,9]. The present results demonstrating bone formation in 3D nHAP-PLGA-collagen scaffolds with kinetics similar to bone regeneration in vivo build upon and expands previous results. Taken together, this combination of biomechanical properties and biochemical signaling is unique; nHAP-PLGA-collagen has superior properties to promote bone regeneration compared to biomaterials lacking one of these components (Table 1). The experimental evidence that nHAP-PLGA-collagen delivers fundamental biological signaling to promote hMSC differentiation into bone tissue while providing a mechanically sound support during osteogenesis inspires further preclinical animal studies to assess the efficacy of nHAP-PLGA-collagen for bone regeneration applications in vivo.

Table 1

Comparison between mechanical and biological properties of different component and combinations

	Mechanical strength in dry condition	Mechanical strength in cell culture condition	Cell viability and proliferation	Osteogenic stimulation
PLGA	High (62)	Low (9)	Low	Low
Collagen	Low (54)	Low (54)	High	Moderate
nHAP-PLGA	High (9)	Moderate (9)	Low	High
nHAP-PLGA-collagen	Moderate (9)	High (9)	High (Fig. 1A)	High (Figs 4 and 5)

While the composition tested in present studies (1.4wt% nHAP, 84.3wt% PLGA and 14.3wt% collagen) demonstrates a good balance between mechanical strength and biological activity for bone regeneration, the mechanical properties, degradation kinetics and osteoinductive properties of nHAP-PLGA-collagen can be adjusted by varying the ratios of the different components comprising the complex biomaterial at the synthesis step. For example, increasing the ratio of hydroxyapatite in nHAP-PLGA-collagen will enhance osteointegration [1,57] as well as increase the stiffness of the biomaterial [17]. Increasing the proportion of collagen increases the elasticity of the scaffold [58,59] and enhances signaling for cell attachment and proliferation [15]. Similarly, changing the ratio of lactide to glycolide monomer in the PLGA [60,61], or including L-lactide instead of D,L-lactide [62] alters the initial mechanical properties and degradation rate. This ability to independently adjust the mechanical properties, bioactivity and degradation rate makes nHAP-PLGA-collagen suitable for a wide range of applications. Moreover, this platform technology can be expanded to covalently attach other proteins, ECM molecules and growth factors with synthetic and natural polymers to adjust the material properties and biochemical signaling based on requirement for a wider range of applications in tissue engineering and regenerative medicine.

The required mechanical properties of a bone biomaterial vary broadly with the size, density, loading direction and anatomical location of the injury [4,63]. The ability to modulate the biomechanical properties of nHAP-PLGA-collagen is unique and makes this biomaterial particularly suitable for regeneration and repair of bone defects for site specific and patient specific applications. This flexibility in adjusting bioactivity and material properties suggests that nHAP-PLGA-collagen can be designed for direct implantation at the injury site with mechanical strength for immediate load bearing and biological activity to regenerate bone in situ at a rate that matches nHAP-PLGA-collagen degradation. Alternatively, nHAP-PLGA-collagen can be seeded with patient stem cells in vitro and the subsequent bone formed in a bioreactor can be implanted at the site of injury to integrate with surrounding bone. With this approach, the cells can be harvested from the patient to reduce the likelihood of immune rejection and transmission of infections. While additional testing is necessary, results from this study along with the possibility to tailor the biological and mechanical properties for a specific anatomical site or bone-repair application suggests that nHAP-PLGA-collagen is a promising candidate for a bone regeneration biomaterial. This platform technology for fabricating covalently linked multicomponent biomaterials can be extended to incorporate other natural polymers, proteins and growth factors to synthesize biomaterials for broader applications in tissue regeneration.

Footnotes

Acknowledgements

The authors would like to thank Dezhi Wang, MD for immunohistochemical staining and the use of the facility at UAB Center for Metabolic Bone Diseases (CMBD) core. The authors are grateful to Drs. Hamid Garmestani from Georgia Tech, Paul D. Eleazer and Eugenia Kharlampieva from UAB for their valuable discussions.

Conflict of interest

The authors have no conflicts of interests.

Ethical statement

There are no animal experiments carried out for this article.

References

Calori

Mazza

Colombo

and Ripamonti

, The use of bone-graft substitutes in large bone defects: Any specific needs?, Injury 42 (2011), S56–S63. doi:10.1016/j.injury.2011.06.011.

Lian

Guo

Cui

Qiu

and Xu

, The mineralized collagen for the reconstruction of intra-articular calcaneal fractures with trabecular defects, Biomatter. 3(4) (2013), e27250. doi:10.4161/biom.27250.

Polo-Corrales

Latorre-Esteves

and Ramirez-Vick

J.E.

, Scaffold design for bone regeneration, Journal of Nanoscience and Nanotechnology 14(1) (2014), 15–56. doi:10.1166/jnn.2014.9127.

Amini

A.R.

Laurencin

C.T.

and Nukavarapu

S.P.

, Bone tissue engineering: Recent advances and challenges, Crit Rev Biomed Eng. 40(5) (2012), 363–408. doi:10.1615/CritRevBiomedEng.v40.i5.10.

Nandi

S.K.

Kundu

Ghosh

S.K.

D.K.

and Basu

, Efficacy of nano-hydroxyapatite prepared by an aqueous solution combustion technique in healing bone defects of goat, Journal of Veterinary Science 9(2) (2008), 183–191. doi:10.4142/jvs.2008.9.2.183.

Pilia

Guda

and Appleford

, Development of composite scaffolds for load-bearing segmental bone defects, Biomed Res Int. 2013 (2013), 458253.

Dimitriou

Jones

McGonagle

and Giannoudis

P.V.

, Bone regeneration: Current concepts and future directions, BMC Med. 9 (2011), 66. doi:10.1186/1741-7015-9-66.

Bhuiyan

Jablonsky

M.J.

Kolesov

Middleton

Wick

T.M.

and Tannenbaum

, Novel synthesis and characterization of a collagen-based biopolymer initiated by hydroxyapatite nanoparticles, Acta Biomaterialia 15 (2015), 181–190. doi:10.1016/j.actbio.2014.11.044.

Bhuiyan

D.B.

Middleton

J.C.

Tannenbaum

and Wick

T.M.

, Mechanical properties and osteogenic potential of hydroxyapatite-PLGA-collagen biomaterial for bone regeneration, Journal of Biomaterials Science, Polymer Edition 27(11) (2016), 1139–1154. doi:10.1080/09205063.2016.1184121.

10.

Hong

Z.K.

Qiu

X.Y.

Sun

J.R.

Deng

M.X.

Chen

X.S.

and Jing

X.B.

, Grafting polymerization of L-lactide on the surface of hydroxyapatite nano-crystals, Polymer. 45(19) (2004), 6699–6706. doi:10.1016/j.polymer.2004.07.036.

11.

Furukawa

Matsusue

Yasunaga

Shikinami

Okuno

and Nakamura

, Biodegradation behavior of ultra-high-strength hydroxyapatite/poly (L-lactide) composite rods for internal fixation of bone fractures, Biomaterials 21(9) (2000), 889–898. doi:10.1016/S0142-9612(99)00232-X.

12.

Hasegawa

Ishii

Tamura

Furukawa

Neo

Matsusue

et al., A 5–7 year in vivo study of high-strength hydroxyapatite/poly (L-lactide) composite rods for the internal fixation of bone fractures, Biomaterials 27(8) (2006), 1327–1332. doi:10.1016/j.biomaterials.2005.09.003.

13.

Shikinami

and Okuno

, Bioresorbable devices made of forged composites of hydroxyapatite (HA) particles and poly-L-lactide (PLLA): Part I. basic characteristics, Biomaterials. 20(9) (1999), 859–877. doi:10.1016/S0142-9612(98)00241-5.

14.

Kruger

T.E.

Miller

A.H.

and Wang

, Collagen scaffolds in bone sialoprotein-mediated bone regeneration, The Scientific World Journal 2013 (2013), 6.

15.

Ferreira

A.M.

Gentile

Chiono

and Ciardelli

, Collagen for bone tissue regeneration, Acta Biomaterialia 8(9) (2012), 3191–3200. doi:10.1016/j.actbio.2012.06.014.

16.

Anderson

J.M.

Vines

J.B.

Patterson

J.L.

Chen

Javed

and Jun

H.-W.

, Osteogenic differentiation of human mesenchymal stem cells synergistically enhanced by biomimetic peptide amphiphiles combined with conditioned medium, Acta Biomaterialia 7(2) (2011), 675–682. doi:10.1016/j.actbio.2010.08.016.

17.

Jaiswal

Haynesworth

S.E.

Caplan

A.I.

and Bruder

S.P.

, Osteogenic differentiation of purified, culture-expanded human mesenchymal stem cells in vitro, Journal of Cellular Biochemistry 64(2) (1997), 295–312. doi:10.1002/(SICI)1097-4644(199702)64:2<295::AID-JCB12>3.0.CO;2-I.

18.

Hauschka

S.D.

and Konigsberg

I.R.

, The influence of collagen on the development of muscle clones, Proceedings of the National Academy of Sciences of the United States of America 55(1) (1966), 119. doi:10.1073/pnas.55.1.119.

19.

Salomon

D.S.

Liotta

L.A.

and Kidwell

W.R.

, Differential response to growth factor by rat mammary epithelium plated on different collagen substrata in serum-free medium, Proceedings of the National Academy of Sciences 78(1) (1981), 382–386. doi:10.1073/pnas.78.1.382.

20.

Aznavoorian

Moore

B.A.

Alexander-Lister

L.D.

Hallit

S.L.

Windsor

L.J.

and Engler

J.A.

, Membrane type I-matrix metalloproteinase-mediated degradation of type I collagen by oral squamous cell carcinoma cells, Cancer Research 61(16) (2001), 6264–6275.

21.

Declercq

H.A.

Verbeeck

R.M.

De Ridder

L.I.

Schacht

E.H.

and Cornelissen

M.J.

, Calcification as an indicator of osteoinductive capacity of biomaterials in osteoblastic cell cultures, Biomaterials 26(24) (2005), 4964–4974. doi:10.1016/j.biomaterials.2005.01.025.

22.

Tenenbaum

and Heersche

, Differentiation of osteoblasts and formation of mineralized bone in vitro, Calcified Tissue International 34(1) (1982), 76–79. doi:10.1007/BF02411212.

23.

Mikos

A.G.

Thorsen

A.J.

Czerwonka

L.A.

Bao

Langer

Winslow

D.N.

et al., Preparation and characterization of poly (L-lactic acid) foams, Polymer. 35(5) (1994), 1068–1077. doi:10.1016/0032-3861(94)90953-9.

24.

Peter

S.J.

Lyman

M.D.

Lai

H.-L.

Leite

S.M.

Tamada

J.A.

et al., In vitro and in vivo degradation of porous poly(dl-lactic-co-glycolic acid) foams, Biomaterials 21(18) (2000), 1837–1845. doi:10.1016/S0142-9612(00)00047-8.

25.

Skaer

, Chemical cryoprotection for structural studies, Journal of Microscopy 125(2) (1982), 137–147. doi:10.1111/j.1365-2818.1982.tb00331.x.

26.

Schofield

B.H.

Williams

B.R.

and Doty

S.B.

, Alcian blue staining of cartilage for electron microscopy. Application of the critical electrolyte concentration principle, The Histochemical Journal 7(2) (1975), 139–149. doi:10.1007/BF01004558.

27.

Carmona-Moran

and Wick

, Transient growth factor stimulation improves chondrogenesis in static culture and under dynamic conditions in a novel shear and perfusion bioreactor, Cellular and Molecular Bioengineering 8(2) (2015), 267–277. doi:10.1007/s12195-015-0387-6.

28.

Rudolf

and Curcio

C.A.

, Esterified cholesterol is highly localized to Bruch’s membrane, as revealed by lipid histochemistry in wholemounts of human choroid, Journal of Histochemistry and Cytochemistry 57(8) (2009), 731–739. doi:10.1369/jhc.2009.953448.

29.

Wang

Jiang

Proctor

McManaman

J.L.

Lucia

et al., Regulation of renal lipid metabolism, lipid accumulation, and glomerulosclerosis in FVBdb/db mice with type 2 diabetes, Diabetes 54(8) (2005), 2328–2335. doi:10.2337/diabetes.54.8.2328.

30.

Sainte-Marie

, A paraffin embedding technique for studies employing immunofluorescence, Journal of Histochemistry & Cytochemistry 10(3) (1962), 250–256. doi:10.1177/10.3.250.

31.

Gubash

S.M.

and Bennett

E.E.

, Use of terpene-based solvents (Hemo-De, Histoclear, and Shandon and BDH xylene substitutes) in place of xylene in the Ehrlich indole test, Journal of Clinical Microbiology 27(9) (1989), 2136–2137.

32.

Allori

A.C.

Sailon

A.M.

and Warren

S.M.

, Biological basis of bone formation, remodeling, and repair – Part II: Extracellular matrix, Tissue Engineering Part B: Reviews 14(3) (2008), 275–283.

33.

Tsai

K.-S.

Kao

S.-Y.

Wang

C.-Y.

Wang

Y.-J.

Wang

J.-P.

and Hung

S.-C.

, Type I collagen promotes proliferation and osteogenesis of human mesenchymal stem cells via activation of ERK and Akt pathways, Journal of Biomedical Materials Research Part A 94A(3) (2010), 673–682.

34.

Somaiah

Kumar

Mawrie

Sharma

Patil

S.D.

Bhattacharyya

et al., Collagen promotes higher adhesion, survival and proliferation of mesenchymal stem cells, PloS One 10(12) (2015), e0145068. doi:10.1371/journal.pone.0145068.

35.

Prins

H.J.

Braat

A.K.

Gawlitta

Dhert

W.J.

Egan

D.A.

Tijssen-Slump

et al., In vitro induction of alkaline phosphatase levels predicts in vivo bone forming capacity of human bone marrow stromal cells, Stem Cell Res 12(2) (2014), 428–440. doi:10.1016/j.scr.2013.12.001.

36.

Genge

B.R.

Sauer

G.R.

McLean

F.M.

and Wuthier

, Correlation between loss of alkaline phosphatase activity and accumulation of calcium during matrix vesicle-mediated mineralization, Journal of Biological Chemistry 263(34) (1988), 18513–18519.

37.

Arpornmaeklong

Brown

S.E.

Wang

and Krebsbach

P.H.

, Phenotypic characterization, osteoblastic differentiation, and bone regeneration capacity of human embryonic stem cell–derived mesenchymal stem cells, Stem Cells and Development 18(7) (2009), 955–968. doi:10.1089/scd.2008.0310.

38.

Santoni

B.G.

Pluhar

G.E.

Motta

and Wheeler

D.L.

, Hollow calcium phosphate microcarriers for bone regeneration: In vitro osteoproduction and ex vivo mechanical assessment, Biomedical Materials and Engineering 17(5) (2007), 277.

39.

Birmingham

Niebur

and McHugh

, Osteogenic differentiation of mesenchymal stem cells is regulated by osteocyte and osteoblast cells in a simplified bone niche, European Cells and Materials 23 (2012), 13–27. doi:10.22203/eCM.v023a02.

40.

Shafiee

Seyedjafari

Soleimani

Ahmadbeigi

Dinarvand

and Ghaemi

, A comparison between osteogenic differentiation of human unrestricted somatic stem cells and mesenchymal stem cells from bone marrow and adipose tissue, Biotechnology Letters 33(6) (2011), 1257–1264. doi:10.1007/s10529-011-0541-8.

41.

Vines

J.B.

Lim

D.-J.

Anderson

J.M.

and Jun

H.-W.

, Hydroxyapatite nanoparticle reinforced peptide amphiphile nanomatrix enhances the osteogenic differentiation of mesenchymal stem cells by compositional ratios, Acta Biomaterialia 8(11) (2012), 4053–4063. doi:10.1016/j.actbio.2012.07.024.

42.

Davies

J.E.

Matta

Mendes

V.C.

and de Carvalho

P.S.

, Development, characterization and clinical use of a biodegradable composite scaffold for bone engineering in oro-maxillo-facial surgery, Organogenesis 6(3) (2010), 161–166. doi:10.4161/org.6.3.12392.

43.

Karageorgiou

and Kaplan

, Porosity of 3D biomaterial scaffolds and osteogenesis, Biomaterials 26(27) (2005), 5474–5491. doi:10.1016/j.biomaterials.2005.02.002.

44.

Stevens

M.M.

and George

J.H.

, Exploring and engineering the cell surface interface, Science 310(5751) (2005), 1135–1138. doi:10.1126/science.1106587.

45.

Aronow

M.A.

Gerstenfeld

L.C.

Owen

T.A.

Tassinari

M.S.

Stein

G.S.

and Lian

J.B.

, Factors that promote progressive development of the osteoblast phenotype in cultured fetal rat calvaria cells, Journal of Cellular Physiology 143(2) (1990), 213–221. doi:10.1002/jcp.1041430203.

46.

Lian

J.B.

and Stein

G.S.

, Development of the osteoblast phenotype: Molecular mechanisms mediating osteoblast growth and differentiation, The Iowa Orthopaedic Journal 15 (1995), 118–140.

47.

Kulterer

Friedl

Jandrositz

Sanchez-Cabo

Prokesch

Paar

et al., Gene expression profiling of human mesenchymal stem cells derived from bone marrow during expansion and osteoblast differentiation, BMC Genomics 8 (2007), 70. doi:10.1186/1471-2164-8-70.

48.

Pockwinse

S.M.

Wilming

L.G.

Conlon

D.M.

Stein

G.S.

and Lian

J.B.

, Expression of cell growth and bone specific genes at single cell resolution during development of bone tissue-like organization in primary osteoblast cultures, Journal of Cellular Biochemistry 49(3) (1992), 310–323. doi:10.1002/jcb.240490315.

49.

Morrison

S.J.

Shah

N.M.

and Anderson

D.J.

, Regulatory mechanisms in stem cell biology, Cell. 88(3) (1997), 287–298. doi:10.1016/S0092-8674(00)81867-X.

50.

Gregory

C.A.

Prockop

D.J.

and Spees

J.L.

, Non-hematopoietic bone marrow stem cells: Molecular control of expansion and differentiation, Experimental Cell Research 306(2) (2005), 330–335. doi:10.1016/j.yexcr.2005.03.018.

51.

Crowder Spencer

Leonardo

Whittaker

Papathanasiou

and Stevens Molly

, Material cues as potent regulators of epigenetics and stem cell function, Cell Stem Cell 18(1) (2016), 39–52. doi:10.1016/j.stem.2015.12.012.

52.

Lee

Cho

H.-Y.

Bui

H.T.T.

and Kang

, The osteogenic or adipogenic lineage commitment of human mesenchymal stem cells is determined by protein kinase C delta, BMC Cell Biology 15(1) (2014), 1–12. doi:10.1186/s12860-014-0042-4.

53.

Becerra-Bayona

Guiza-Arguello

Munoz-Pinto

D.J.

and Hahn

M.S.

, Influence of select extracellular matrix proteins on mesenchymal stem cell osteogenic commitment in three-dimensional contexts, Acta Biomaterialia 8(12) (2012), 4397–4404. doi:10.1016/j.actbio.2012.07.048.

54.

Roeder

B.A.

Kokini

Sturgis

J.E.

Robinson

J.P.

and Voytik-Harbin

S.L.

, Tensile mechanical properties of three-dimensional type I collagen extracellular matrices with varied microstructure, J Biomech Eng. 124(2) (2002), 214–222. doi:10.1115/1.1449904.

55.

Mima

Miya

Iwamoto

and Yoshikawa

, Highly deacetylated chitosan and its properties, Journal of Applied Polymer Science 28(6) (1983), 1909–1917. doi:10.1002/app.1983.070280607.

56.

Zhang

Ouyang

Lim

C.T.

Ramakrishna

and Huang

Z.-M.

, Electrospinning of gelatin fibers and gelatin/PCL composite fibrous scaffolds, Journal of Biomedical Materials Research Part B: Applied Biomaterials 72B(1) (2005), 156–165. doi:10.1002/jbm.b.30128.

57.

Moore

W.R.

Graves

S.E.

and Bain

G.I.

, Synthetic bone graft substitutes, ANZ Journal of Surgery 71(6) (2001), 354–361. doi:10.1046/j.1440-1622.2001.02128.x.

58.

Gastaldi

Parisi

Lucchini

Contro

Bignozzi

Ginestra

P.S.

et al., A predictive model for the elastic properties of a collagen-hydroxyapatite porous scaffold for multi-layer osteochondral substitutes, International Journal of Applied Mechanics 07(04) (2015), 1550063. doi:10.1142/S1758825115500635.

59.

Burstein

Zika

Heiple

and Klein

, Contribution of collagen and mineral to the elastic-plastic properties of bone, The Journal of Bone & Joint Surgery 57(7) (1975), 956–961. doi:10.2106/00004623-197557070-00013.

60.

Makadia

H.K.

and Siegel

S.J.

, Poly Lactic-co-Glycolic Acid (PLGA) as biodegradable controlled drug delivery carrier, Polymers 3(3) (2011), 1377–1397. doi:10.3390/polym3031377.

61.

Gentile

Chiono

Carmagnola

and An

H.P.

, Overview of poly(lactic-co-glycolic) acid (PLGA)-based biomaterials for bone tissue engineering, International Journal of Molecular Sciences 15(3) (2014), 3640. doi:10.3390/ijms15033640.

62.

Ferguson

Wahl

and Gogolewski

, Enhancement of the mechanical properties of polylactides by solid-state extrusion. II. Poly (l-lactide), poly (l/d-lactide), and poly (l/dl-lactide), Journal of Biomedical Materials Research 30(4) (1996), 543–551.

63.

Pilia

Guda

and Appleford

, Development of composite scaffolds for load-bearing segmental bone defects, BioMed Research International 2013 (2013), 1–15.