Bone is an organic-inorganic composite with the ability to regenerate itself. Thus, several studies based on artificial organic-inorganic interface sciences have been tried to develop capable materials for effective regenerative bone tissues. Hydroxyapatite nanoparticles (HAp NPs) have extensively been researched in bone tissue engineering due to the compositional and shape similarity to the mineral bone and excellent biocompatibility. However, HAp alone has low mechanical strength, which limits its applications. Therefore, HAp NPs have been deposited on the biocompatible polymer matrix, obtaining composites with the enhanced mechanical, thermal, and rheological properties and with higher biocompatibility and bioactivity. For developing new biomedical applications, polymer-HAp interfacial interactions that provide biofusion should be investigated. This paper reviewed common coating techniques for obtaining HAp NPs/polymer fusion interfaces as well as
Human bone is an organic-inorganic composite having the components of collagen fibrils containing embedded and well-disposed, nanocrystalline hydroxyapatite (HAp) with 25–50 nm length and rod-like shape [
Bioceramics are biocompatible ceramics that have been used for biomedical applications in both crystalline and amorphous forms [
Bioinert bioceramics (e.g., titania (TiO2) [
Bioactive bioceramics (e.g., hydroxyapatite (HAp), tricalcium phosphate, biphasic calcium phosphates, bioactive glass, and some glass-ceramic formulations [
Resorbable or biodegradable bioceramics (e.g., calcium phosphates like tricalcium phosphate (TCP) and carbon-contained HAp (CHAp) [
Nanotechnology has allowed the preparation of nanostructured biomaterials, giving way to the third generation of bioceramics, focused on the enhanced bioactivity and the initial physiological trace inducing an enhanced cell to respond at the molecular level in order to regenerate tissues due to their similarity to the inorganic component of human bone tissues [
In this paper, HAp nanoparticles (NPs) have been deposited on the biocompatible polymer matrix, obtaining composites with enhanced mechanical, thermal, and rheological properties and with higher biocompatibility and bioactivity. For developing new biomedical applications, polymer-HAp interfacial interactions that provide biofusion should be investigated. This paper reviewed common coating techniques for obtaining HAp NPs/polymer fusion interfaces as well as
The inorganic component of bone and teeth is nanocrystalline HAp, which provides the toughness and ability to withstand pressure. This calcium phosphate is stacked and aligned with the organic matrix formed by collagen fibers, glycoproteins, and mucopolysaccharides for conferring the elasticity and resistance to the structure [
Despite the great variety of compounds that can be prepared by substituting elements at the
(a) Crystal structure of HAp in the top view along with the
The properties of HAp can significantly affect the particle composition, assembly, size, and morphology. Nanosized HAp NPs (20–80 nm) have shown to be more efficient in osteoblast adhesion and proliferation and improved mineralization [
The obtained sizes and shapes of HAp depend strongly on the synthetic route as well as synthetic parameters [
TEM images of (a) general HAp (HAp-40-Aft) NPs, (b) cationic surfactant-assisted HAp (CTAB/HAp-40-Aft) NPs, (c) zinc ion-doped HAp (0.5-Zn : HAp) NPs, and (d) zinc ion-doped CHAp (0.5-Zn : HAp) NPs. Reprinted with permission from [
HAp-based nanophase materials offer excellent biocompatibility, biodegradability, and osteoconductive and osteoinductive properties [
Illustration of the various biomedical applications of HAp NPs.
Several forms, shapes, and sizes of HAp have been synthesized and investigated; however, nanosized HAp exhibits enhanced bioactivity, biocompatibility, mechanical properties, and higher resorbability as compared with microsized HAp [
Studies on polymers applied to the medical field began in the 20th century; nonetheless, at the end of the 1950s, the use of polymers in medicine and medical applications was intensified because of their good biocompatibility, low toxicity, bioinert nature, and good mechanical properties such as strength, abrasion resistance, and flexibility [
Features of representative biomedical polymers used
Polymer | Basic properties | Possible biomedical applications | References |
---|---|---|---|
Collagen | Biocompatible, biodegradable, great tensile strength, and weak antigenicity. The isoelectric point is 8.26. Young’s modulus in the range from 3.7 to 11.5 GPa | Wound healing, tissue engineering, hemostatic agent, bone grafts | [ |
Chitin | Nontoxic, biodegradable, biocompatible, and antimicrobial and hydrating agent. Highly hydrophobic, insoluble in water and even most organic solvents | Tissue engineering scaffolds, drug delivery, wound dressings, antibacterial coatings, and sensors | [ |
Chitosan (CS) | Nontoxic, biodegradable, and biocompatible polymer. Tensile strength is 0.0650 MPa, Young’s modulus in the range from 0.00443 to 0.0236 MPa; it is degraded after 220°C | Tissue engineering scaffolds, bone regeneration, angiogenesis, and wound healing | [ |
Poly( |
Biodegradable, amphiphilic, flexible, nontoxic, and biocompatible polyester with melting point of around 60°C and a glass transition temperature of about −60°C | Tissue engineering, long-term implantable devices, drug delivery systems, microencapsulation, and scaffold for tissue repair | [ |
Polyurethane elastomers (eLPU) | Biocompatible, biodegradable and with tailorable chemical and physical forms. Glass transition temperature about −73 to −23°C, Young’s modulus in the range from 0.002 to 0.003 GPa, tensile strength, and yield stress in a range from 25 to 51 MPa | Used in catheters, drug delivery vehicles, prosthetic implants, surgical dressings/pressure sensitive adhesives, tissue engineering scaffolds, and cardiac patches | [ |
Poly(amide) (PA) | High crystallinity, good mechanical properties including high tensile strength, high flexibility, good resilience, low rates of biodegradation, very high tenacity, and excellent sliding characteristics and wear resistance. Conductivity: 10−12 S/m, melting point: 190–35°C, thermal conductivity: 0.25 W/(m·K) | Used as suture material, ligament and tendon repair, balloon of catheters, and dialysis membranes | [ |
Polyether ether ketone (PEEK) | Semicrystalline, excellent mechanical, very stable, and chemical resistance properties. Glass temperature: 143°C, Thermal conductivity: 0.25 W/(m·K), melting point: 343°C, Young’s modulus: 3.6 GPa, tensile strength: 90–100 MPa | Orthopedic applications or inner lining of catheters | [ |
Poly(ethylene terephthalate) (PET) | Strong and impact-resistant, excellent water and moisture barrier material, biostable, insoluble in water, melting point: >250°C, glass transition temperature: 67–81°C, Young’s modulus: 2800–3100 MPa, tensile strength: 55–75 MPa thermal conductivity: 0.15 to 0.24 W/m.K | Used for membranes, vascular grafts, surgical meshes, and ligament and tendon repair | [ |
Poly(ethylene glycol) (PEG) | Nontoxic, nonimmunogenic, nonantigenic, hydrophilic, bioresorbable, and biocompatible polymer. Flash point: 182–287°C. | Used as antifouling coating on catheters, hydrogel or as pore former in dialysis membranes and drug delivery systems | [ |
Poly(dimethylsiloxane) (PDMS) | Hydrophobic, stable, biocompatible, bioinert, flexible, and soft rubbery behavior | Used for catheters, nucleus pulposus substitute, plastic surgery, intraocular lenses, glaucoma drainage devices, and dialysis membranes | [ |
Polyglycolic acid (PGA) | High crystallinity (45–55%), high tensile modulus, poor solubility in organic solvents. Excellent fiber forming ability. High rate of degradation and acidic degradation products. Glass transition temperature: 35–40°C and melting point 235–230°C, tensile strength: 340–920 MPa, Young’s modulus 7–14 GPa | Regenerative biological tissue, bone internal fixation devices, tissue engineering scaffolds, suture anchors, meniscus repair, medical devices, and drug delivery | [ |
Poly(L-lactic acid) (PLLA) | Degradable, good tensile strength. glass transition temperature: 50–60°C and melting point 170–190°C, tensile strength: 870–2300 MPa, Young’s modulus 10–16 GPa | Orthopedic fixation tools, ligament and tendon repair, and vascular stents | [ |
The chemical compositions and structures of the polymer surfaces will determine the interfacial interactions that take place between the biological media (such as proteins, cells, and tissues) and the polymer substrates [
One of the most used polymers in biomedical applications is polydimethylsiloxane (PDMS), due to the good biocompatibility, low toxicity, high flexibility, good thermal and oxidative stability, low modulus, antiadhesive nature, soft and rubbery behavior, bioinertness, transparency in the visible region, and control of free volume with the aim of accommodation of metal nanoparticles [
Dependence of the cross-linking degree on the incorporation of gold (Au) nanoparticles in the PDMS film surface modification. (a) General reaction between the cross-linker and liquid PDMS, (b) UV-visible absorption spectra, and (c) photographs of the Au-PDMS composite films containing (i) 2, (ii) 4, (iii) 6, (iv) 8, and (v) 10 wt% of the cross-linker to liquid PDMS. Reprinted with permission from [
The surface structure of the biocompatible polymer is able to enhance protein adsorption, and then the cells interact with the proteins, leading to the cells forming tissues on the biomedical polymer. In this manner, it is desirable to tailor the surface of the biomedical polymer in order to provide a biocompatible physicochemical environment to guide the cells to form tissues [
Two commonly used strategies to functionalize polymers can be raised: (1) functional groups are introduced into polymer monomers. Although hydrophilic co-monomers can be inserted into the prepolymerization system by chemical functionalization, these monomers can alter the bulk properties, which is not desirable for biomedical applications [
By plasma-enhanced CVD (PECVD), different kinds of polymeric surfaces can form different types of thin films depending on inorganic coatings such as carbon nanotubes [
Surface-functionalized biomedical polymers are expected to show good potential properties for blood-contacting devices [
For tissue engineering, cell compatibility is necessary. Our group studied the effect of vitronectin and
Effects of plasma treatment of PDMS surfaces. (a) Schematic representation of the changes from hydrophobic to hydrophilic PDMS surface, and the corroboration by a static contact angle with 2 mL water droplets. (b) Schematic results of the adhesion of human-induced pluripotent stem cells (hiPSCs) to plasma-patterned polydimethylsiloxane coated with vitronectin (Vn) and
Artificial bone is one of the most transplanted tissues. For this reason, the interest to develop new and more biocompatible inorganic-organic composites is increasing [
The objective of the incorporation of HAp into the biocompatible polymer matrix enhances the mechanical strength and provides the topographic features to improve the integrity of implants and the surrounding bone and to stimulate bone tissue ingrowth [
(a) Scheme of the electrophoretic deposition (EFD) procedure with a direct current (DC) of 100 V, (b) deposition mechanism in ethanol, and (c) resultant electrophoretic deposition system at the monolayer, after ultrasonification. Reprinted with permission from [
Silicone surface modification technique through the deposition of zinc- (Zn-) substituted HAp by EFD to achieve for providing bioactive properties to the films. Photographs of the PDMS substrates (a) before and (b) after the EPD and (c) after ultrasonication (US), indicating the maintenance of a flexible state in folding back. (d) AFM topographic images of 0.5-Zn : HAp nanocrystals electrophoretically prepared at coating voltages of 100 V in ethanol and micrographs of the biocompatible properties based on fibroblast ingrowth on the Zn-substituted HAp with the different initial Zn concentrations of (e) 0.0, (f) 2.5, (g) 5.0, and (h) 10 mol% at the culture time of 72 h. Reprinted with permission from [
Biomineralization (or biological mineralization) is a well-regulated process, which is responsible for the controlled formation of inorganic materials from aqueous solution in living organisms [
Scheme of the biomineralization process. (a) Ions in solution, (b) nucleation, (c) aggregation, (d) amorphous calcium phosphate—the circle shows the Posner cluster unit [Ca9 (PO4)6] projected on the
The new technology called “bioinspired growth” has been sought to emulate the natural biomineralization process in order to obtain bioactivity and mechanical properties, which can improve the biointeractions with the human bone [
Scheme of the biomineralization process to promote biological apatite (BAp) growth from a simulated body fluid (SBF) by three different processes: process 1—the bare substrates induced BAp growth by immersion in SBF; process 2—the fetal bovine serum (FBS) proteins preadsorbed on the substrates showed slight BAp growth, indicating a significant inhibition of the BAp growth; and process 3—the BAp coating technique on tissue culture poly(styrene) through the film formation by the hybridization of BAp with the L-
It has been observed that incorporation of HAp into the polymer matrix can enhance the mechanical properties, increase the roughness, and produce a topography that allows mimicking the nanostructure of the bone [
The process of bone tissue formation is called osteogenesis and is carried out by two ossification mechanisms: intramembranous (IM) and endochondral (EC) [
For porous composite structures, the study of the interactions between the HAp/bioinert polymer and the cells is carried out using
The initial attachment of human osteoblasts (HOBs) on poly(
Study of hepatocyte cell aggregation and adhesion at HAp NPs covered with SU-8 polymer micropatterns by nano/microfabrication techniques. (a) Schematic illustration of the micropatterning process to obtain SU-8/HAp nanocrystals/Au/Ti/silica substrate and the subsequent cells adhered on HAp surface, (b) optical microscopic image of a cell, (c) 3D graphics image of the cell, (d) AFM topographic, and (e) phase shift images of the cell and HAp nanocrystalline surfaces. Reprinted with permission from [
Cell adhesion, proliferation, migration, differentiation, and survival can be modulated by ECM proteins. The ECM can influence diverse types of cells such as osteoblasts, osteoclasts, osteocytes, and bone lining cells [
Possible schemes of the protein interactions with the hydration layer on the HAp surface during protein adsorption. (a) In the beginning, hydrated fibrinogen (Fgn) interacts with the hydration layer with possible dehydration. Then, Fgn absorbs on HAp. Finally, conformational change from “side-on,” at initial adsorption region, to “end-on” at the saturated adsorption region. (b) The adsorption model of albumin (Ab) could be “side-on” at the initial adsorption region and also at the monolayer. Nonfreezing water could suppress the denaturalization of the adsorbed proteins suggesting fusion interfaces. Reprinted with permission from [
Several HAp/polymer/cell interactions have been studied to improve bone tissue regeneration and to control cell adhesion. HAp which bonded with insulin was incorporated in PLGA to obtain insulin-HAp/PLGA composites.
(a) Scheme of the evaluation of the effect on the interfacial phenomenon during the initial adhesion between the proteins and the osteoblast-like cells by QCM-D and the confocal laser scanning microscope (CLSM). CLSM images of the cells adhered on (b) fetal bovine serum (FBS), (c) FBS-Fgn, (d) FBS-Ab, and (e) FBS-collagen adsorbed on the HAp. Reprinted with permission from [
Figure
(a) Illustration of the conventional successive events on bioceramic surfaces after the implantation into the animal body. (b) Possible illustration of the preferential mobility for cell adhesion by the fusion interfaces between the bioceramics and polymers, exhibiting comfortable viscoelastic and flexible structures by the cells.
HAp NPs, biocompatible polymers, and their composites have been extensively studied in both
The above points have revealed some of the critical events for HAp NPs to stimulate an interface in body fluid. In future, the study on biointeractive interfaces of HAp NPs deposited on the rigid and soft polymers can be useful for biomedical applications. Due to that the HAp/rigid polymer has better support and strength, it can be used for bone repair and replacement in hard tissues, while the HAp/soft polymer can be used in skin tissues that require more flexibility to preserve their functions, as is the case with tendon repair and cartilage replacement, to build blood vessels or long-term catheters. Further researches can bring significant improvements to existing experimental methods to prepare and characterize useful nanobioceramic-polymer fusion interfaces. These studies will lead to a deep understanding of nanobiointerfaces.
To overcome existing scientific challenges, mutual interactions at the nanobiointerfaces should be explored by developing novel detection techniques for biomolecular interactions [
There is no conflict of interest regarding the publication of this paper.
This study was supported by a grant from the Japan Society for the Promotion of Science (JSPS) KAKENHI (Grant-in-Aid for Young Scientists (A), Grant No. 17H04954).