Review: physico-chemical modification as a versatile strategy for the biocompatibility enhancement of biomaterials

Abstract

A biomaterial can be defined as a material intended to interface with biological systems to evaluate, treat, augment or replace any tissue, organ or function of the body. Major problems associated with biomaterials are their properties and biocompatibility, which need to be tackled and resolved before promoting a particular biomaterial to the market or implanting it into a biological system. To enhance the biocompatibility of the biomaterials, several surface modification strategies, such as physico-chemical, mechanical and biological modifications, have been explored. In this review, some recent applications of physico-chemical modification technologies, such as alteration in the structure of a molecule by chemical modification, surface grafting, abrasive blasting and acid etching, surface coatings, heat and steam treatment for medical materials such as polymers, metals, ceramics and nanocomposites are discussed. This article will promote physico-chemical modification as a versatile technology in surface engineering to improve the properties and biocompatibility of medical materials. Furthermore, it will instigate the growth of the biomaterial market with various high quality biomaterials.

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1 Introduction

Biomaterials are used to make medical devices that can replace a part or a function of the body in a safe, reliable, economic and physiologically acceptable manner.¹ The global biomaterials market is expected to record close to a 15% yearly growth for the ten-year period ending 2017, reaching around $84 billion.² Recently, biomaterials have enjoyed widespread application in various biological systems, such as in skeletal (bone plate, total joint replacements), muscular (sutures, muscle stimulator), nervous (cardiac pacemaker, nerve stimulator), endocrine (microencapsulated pancreatic islet cells), reproductive system (augmentation mammoplasty), dental and maxillofacial applications (cosmetic replacements), as well as drug-delivery systems.^3,4 Biomaterials are mainly classified into four main categories, namely, metals, ceramics, polymers and biological substances. The selection of a biomaterial depends on the surrounding environment where it will be implanted. The implanted material should not cause any adverse effects like allergies, inflammation and toxicity, either immediately after surgery or under post-operative conditions.⁵ The first requirement of a biomaterial is that the material must be biocompatible; this means that the organism should not treat it as a foreign object. Biocompatibility is a fundamental property that decides the excellence of a biomaterial and its application in the medical field. The term biocompatibility denotes the ability of a material to perform with an appropriate host response in a specific situation.⁶ Biocompatibility has been discussed in lots of works with escalating curiosity in assessing the characteristics of medical materials and devices and also the responses caused by their components. Biocompatibility covers many aspects of the material, including its physical, mechanical and chemical properties, as well as toxicity, mutagenic and allergenic effects, with the aim that no noteworthy injuries or harmful effects on the biological functions of cells and individuals should take place. In order to gain a better knowledge, readers may refer to the following cited articles.^7,8 Blood compatibility refers to the events taking place within the biological system when the material surface comes into contact with the blood and its components. Blood compatibility is the key property of the implant material, especially for devices that come into contact with blood.⁹ Thus, we can describe blood compatibility as the ability of the material to perform its function in a particular situation without causing any blood-related complications. Whenever the blood comes into contact with implants (biomaterials), it can directly lead to complications, such as the interaction of blood components with surfaces resulting in protein and water adsorption or blood cells interfering with the surface of biomaterial, and these events can lead to haemostasis and coagulation.¹⁰ Until the biocompatibility of a material is confirmed, it must be subjected to various studies ranging from in vitro assays to clinical trials, in the areas of pharmaceutics, biology, chemistry and toxicology to validate its use as a biomaterial. Secondly, the material should allow the biological system to resume its natural functioning. Thirdly, the material should be mechanically sound; for the replacement of load bearing structures, the material should possess an equivalent or greater mechanical stability to ensure high reliability of the graft.¹¹ It is vital to modify the biomaterial surfaces in order to control the subsequent interaction of the implant surface with blood or the biological system and it responses for particular applications.

Surface modification approaches, namely, physico-chemical, mechanical and biological methods, are currently in use. These three categories are further subdivided into sub-categories. In our review we mainly focus on the physico-chemical modification of biomaterials to improve the blood compatibility of the biomaterial. The most important techniques involved in physico-chemical methods are modifying the surface by chemical means, surface grafting, abrasive blasting and acid etching, surface coating, high temperature treatments (thermal treatment, vapour and steam treatment), as shown in Fig. 1.

Fig. 1 Physico-chemical modification technologies and strategies.

2 Surface treatment

The surface characteristics of all types of biomaterials play a key role in determining the biocompatibility of a particular implant. When choosing the biomaterial to communicate with the biological system, the first important criterion to inspect is that of biocompatibility.¹² The biocompatibility and cellular interactions of the biomaterial are based on the surface characteristics. Properties such as surface roughness, hardness, temperature, surface chemistry, surface reactivity (inert or active), wettability and surface charge (surface free energy) are the surface characteristics that play a major function in cell adhesion, cell spreading, cell proliferation and tissue formation.¹³ The response of the host organism in macroscopic, cellular and protein levels to biomaterials is closely associated with the surface properties of the material. The appropriate physico-chemical properties of a biomaterial required for implantation will depend on the corresponding biomedical application and can be modified by performing surface treatment techniques. Surface modification is the process of modifying the surface of a material by changing the physical or chemical characteristics from the ones originally found on the surface of the material, which can be carried out in a nanoscale or bulk level. Nanoscale surface modification is an important technique in the field of nanotechnology, which involves nanofabrication that may modify both the topography and the chemistry of the surface at the nanometric level. In nanoscale modification, the modifications of substrates are carried out in the nano domain. For example, in coating technology, the thickness of the films is in the range of nanometers. The nanomolecular layers are less than 10 nm thick and are used to produce nanoscale modifications.¹⁴ After performing the nanoscale modification, the effect of modification is studied by nanometric analysis, including scanning tunneling microscopy (STM) and atomic force microscopy (AFM), which shows nanolevel pit formations and the average roughness in nanometers. Parameters such as cell adhesion and proliferation involved in the interaction mechanism of the modified surface with biological substances can also be controlled in the nanoscale.¹⁵ The surface coating with nanoparticles decreases the negative properties of the material and enhances the osteoblast cell adhesion and proliferation, as compared with microparticle coatings. It has been noted that the smooth surface on a microlevel surface modification is not necessarily needed to be smooth in nanolevel modification.¹⁶ On the other hand, nano-engineered surfaces can directly influence the biological properties and will be more useful in the applications of tissue regeneration functionalities than bulk surface modification,¹⁷ while in bulk modification, a large surface is subjected to modification techniques in order to modify its bulk properties. During bulk surface modification, the bulk properties, i.e. the properties resulting from the greater number of atoms present in the sample, gets changed. In contrast to nanoscale surface modifications, the objective of bulk modification is to change a wide range of properties, including the mechanical, physical and chemical characteristics.¹⁸ The interactions between blood and a material surface depend on the blood composition, blood flow and the surface characteristics of the implanted material defined by its physico-chemical properties.¹⁹ The modification can be done by different methods with an objective to altering a wide range of surface properties, such as surface roughness, hydrophilicity, surface charge, biocompatibility and reactivity.^20–23 In this review, we will discuss chemical modification, surface grafting, abrasive blasting and acid etching, surface coating, thermal treatment, vapour and steam treatment, i.e. the major techniques for physico-chemical modification to enhance the biocompatibility of the biomaterial.

3 Chemical modification

Chemical modification includes alkali hydrolysis, covalent immobilization and wet chemical methods, although these are only three of the many ways to chemically modify a surface. The surface is first prepared by surface activation, where several functionalities are positioned on the material surface to chemically modify the surface, as studies have shown that chemical modification enhances biocompatibility. The wet chemical method is one of the mostly chosen methods, where the chemical species are dissolved in an organic solution and reactions take place to reduce the hydrophobic nature. The surface stability is higher in chemical modification than in physical adsorption. It also offers better biocompatibility towards cell growth and bodily fluid flow. Thus, chemical modification is a significantly used surface treatment for all types of biomaterials such as polymers, metals, ceramics and nanocomposites.

3.1 Chemical modification of polymers

Lim et al. established that biological progressions, including protein adsorption, cell proliferation, and gene expression, can be restricted to some extent using chemical methods to modify the surface properties of biocompatible materials, leading to controlled surface functionalization of the material.²⁴ It was stated that poly(lactic-co-glycolic acid) can be nanostructured by chemical etching with sodium hydroxide (NaOH), resulting in material surface features that could be used to enhance the activity of various cell types.^25,26 Balakrishnan et al. subjected polyvinyl chloride (PVC) to amination under a concentrated solution of ethylenediamine, followed by PEG treatment. In that study, platelet studies were performed and it was concluded that the platelet adhesion is significantly reduced for the modified PVC compared to a control PVC. The contact angle measurements depicted an increase in the hydrophilicity of the modified polymeric surface.²⁷ Yvette et al. tailored the surface of polyurethanes (PUs) by the covalent attachment of dipyridamole (trademarked as Persantine) to confirm the inhibition of thrombus formation and adherence of the blood platelets upon exposure to human platelet rich plasma (PRP).²⁸ The polyethersulfone (PES) membrane was modified for hydrophilic functionality to depress protein adsorption and platelet adhesion. The activated partial thromboplastin time (APTT) for the modified PES membranes was increased, resulting in enhanced blood compatibility.^29,30 Saravana et al. investigated the blood compatibility of metallocene polyethylene (mPE) after treating the polymer surface with hydrochloric acid (HCl). The contact angle analysis of the treated sample indicated an increase in hydrophilicity. FTIR results showed that there were no notable changes on the functional group, while SEM images of the modified samples proved that the acid-tailored surface was engaged with the formation of pits. Blood coagulation assays like prothrombin time (PT) and activated partial thromboplastin time (APTT) revealed that there was a delay in the clotting on the surface of treated samples. The outcome of a haemolytic assay depicted minor damage to the red blood cells (RBC) compared to the untreated sample. A platelet adhesion assay showed that the number of platelets adhering on the surface of the treated polymer was considerably less than that on the untreated surface.³¹ Surface modification of polyurethane (PU) was performed by blending the sulfonated polyrotaxanes (PRx-SO(3)'s) with a PU solution, followed by solution casting. The incorporation of PRx-SO(3)'s on PU led to the enhancement of hydrophilicity by changing the surface properties of the PU matrix. Thus, surface modification with PRx-SO(3)'s is recommended as valuable for the fabrication of biocompatible medical devices.³² Poly(acrylonitrile-co-maleic acid)s (PANCMAs) were tethered with poly(ethylene glycol)s (PEGs). Chemical modifications on the membrane surface were characterized by Fourier transform infrared spectroscopy and the hydrophilicity and blood compatibility of the PEG-tethered PANCMA membrane were examined by water contact angle, plasma platelets adhesion and cell adhesion measurements. The results revealed that the hydrophilicity of the membrane can be improved significantly, and the protein adsorption, platelets adhesion and macrophage attachment on the membrane surface were clearly suppressed.³³ Chemical modification of the polymers may produce new polymeric materials, which cannot be synthesized by the polymerization of monomers and is considered an advantage of the chemical modification of polymers, even though non-specific interactions also exist.³⁴

3.2 Chemical modification of metals

Hansson et al. treated the titanium material surface with polyethylene glycol (PEG) and proved that after PEG treatment the material has outstanding blood compatibility by providing resistance to the adhesion of small bio-molecules like fibrinogen and cells such as platelets and leukocytes.³⁵ The surface chemical composition was also changed by means of alkali modification to improve the biocompatibility as well as the bioactivity of a titanium (Ti) implant. The chemical modification of the Ti implant with hydrofluoric (HF) acid resulted in the reduction of the hydrocarbon content, which led to an increase in the surface energy and potential for bio-acceptability of the Ti-implant.³⁶ NaOH treatment of the titanium implant resulted in the formation of sodium titanate on the treated surface and the histological assessment and scanning electron microscopy (SEM) observation showed that new bone was formed on the surface of the alkali-modified implants and that the bone grew more rapidly than the unmodified implants.^37,38 The cytotoxicity evaluation of phytic acid-treated WE43 magnesium alloy showed that the biocompatibility of the phytic acid-treated WE43 Mg alloy is much better than blank WE43 magnesium alloy. From the haemolysis test results, we can infer that the modified samples with more Phy–Mg complex will have a better biocompatibility.³⁹ In another study, the pure titanium (Ti) and titanium alloy (Ti–6Al–4V) specimens were implanted into mice with and without any surface treatment. Three months after implantation, the biocompatibilities of the unmodified and modified implants were examined by in vitro and in vivo experiments. The outcome of these experiments showed that the commercial pure Ti and Ti–6Al–4V alloy specimens treated with alkali (KOH) had a better biocompatibility than commercial pure Ti and Ti–6Al–4V alloy specimens without alkali treatment.⁴⁰ Parsapour et al. treated stainless steel with HNO₃, followed by H₂SO₄ and ended the surface treatment with a Nb coating to form a passive layer on the surface. The end product of the acid treated, Nb-coated stainless steel showed improved biocompatibility.⁴¹ It has been shown then that the chemical modification of metals produces desired topographical properties with enhanced biocompatibility, which is directly based on the chemical reagents used for the surface modification for a particular application.

3.3 Chemical modification of ceramics

The surface chemical modification of ceramics to improve the biocompatibility and the influence of chemical treatment on cellular behavior were studied. An in vitro study inspected the effects of the surface chemistry modification of bio-ceramics on human bone-derived cells (HBDCs) and concluded that the surface chemistry affects the cell adhesion,⁴² and that a negative potential was effective in increasing the adhesiveness with increasing wettability, even though living cells have negative charges.⁴³ Al₂O₃ bio-ceramic was implanted with NH₂⁺ ions and it was found that the quantity of amidogen radicals implanted on the ceramic surface was proportional to the dosage of NH₂⁺ ions used during the ion implantation process. In addition, when an implantation power of 100 keV was used, the highest amount of NH₂ radicals were implanted on the Al₂O₃ ceramic surface. The results of the biocompatibility test show that the ceramic surfaces implanted with NH₂⁺ ions have better biocompatibility, compared to the unimplanted Al₂O₃ bio-ceramic surface.⁴⁴ Calcium silicate (CS), a biodegradable ceramic, was chemically modified by partially replacing the calcium sites by strontium. The SEM images of the modified ceramic surface indicate an improved bioactivity, as well as improved biocompatibility of the ceramic material.⁴⁵ The surface of yttria-tetragonal zirconia polycrystal (Y-TZP) was modified by a hydrothermal treatment. The topographies of modified Y-TZP specimens were analyzed by contact angle assay, XRD, FTIR, AFM, and FE-SEM. Then, the RGD-peptide was immobilized on the surface of the Y-TZP by chemical treatment and the resultant surface was analyzed by SEM, FTIR. The results indicate that the cell activity and biocompatibility were better for the RGD-peptide immobilized Y-TZP than that for the unmodified Y-TZP.⁴⁶ From the reported literature, the bioactivity of the chemically modified ceramics was significantly increased and the selection of the appropriate chemical reaction for a particular application remains a challenge.

3.4 Chemical modification of nanocomposites

In the surface modification of nanocomposite scaffolds, gelatin was initially entrapped onto the surface and then heparin was subsequently immobilized on the entrapped gelatin. The surface-modification improved the wettability of scaffolds.⁴⁷ Nano-hydroxyapatite/poly(ε-caprolactone) (PCL) particles were modified with the silane coupling agent, KH-792, and showed a more positive effect on biocompatibility than the control group.⁴⁸ PVA (polyvinyl alcohol)/starch composites were subjected to surface treatment to enhance their biocompatibility. The modified surfaces were studied by FTIR and contact angle measurements. The results of this study concluded that the surface characteristics were based on the type and number of incorporated nanoparticles, as well as on the treatment applied.⁴⁹ Nano-hydroxyapatite (nHA) was wrapped using polypropylene glycol (PPG), and then these nHA particles were successfully introduced on the polyurethane surface. Coagulation assays were performed and displayed the delay in clotting time, while the MTT assay confirmed the biocompatibility of the modified nano-hydroxyapatite (nHA) composite.⁵⁰ Adhikari et al. studied polymer–matrix nanocomposites based on poly(lactic-co-glycolic) acid (PLGA) and graphene platelets (GNPs). Their biocompatibility was then examined, and suggested that the PLGA/GNP nanocomposites showed better biocompatibility for cell growth with/without graphenes functionalization compared to pure PLGA.⁵¹ The nanomaterial was surface modified with polydopamine (PDA) in a controlled manner compared to the water-phase polymerization. A PDA-shelled nanocomposite showed reduced toxicity and enhanced biocompatibility.⁵² Fe₃O₄ nanoparticles were modified by thiodiglycolic acid (TDGA) and used in the preparation of magnetite nanoparticles with improved mechanical properties; a study with fibroblast cell interaction showed that the modified surface had good biocompatibility.⁵³ The surfaces of the nanocapsules were modified with polyethylene oxide (PEO) and succinic anhydride, and biomedical tests such as haemolysis and thromboelastography (TEG) were conducted over the surface modified nanocapsules. The outcome of these experiments showed that the PEO surface modification greatly reduced the damaging interactions of the nanocapsules with red blood cells (RBCs) and platelets.⁵⁴ The surface of the micron-sized hydroxyapatite (HA) particles was modified by the in situ polymerization of styrene (St), then compounded with high impact polystyrene (HIPS). The surface-treated HA particles displayed improved biocompatibility.⁵⁵ The chemical modification of nanocomposites adds an advantage of agglomeration reduction, as well at the effects of the modification depending on the chemicals used.⁵⁶

4 Surface grafting

Surface grafting refers to the addition of polymer chains to the surface to change the surface properties. A thin film can be formed on the material surface through spin casting, precipitation, the Langmuir–Blodget technique, polymer adsorption and chemical grafting. Among these techniques, chemical grafting offers more advantages than the other methods because of the ease and controllable addition of a number of polymer chains on the same material surface with high surface density, precise localization and long stability of the grafted layers. Surface grafting offers existing materials new functionalities, such as hydrophilicity, adhesion, biocompatibility and anti-fogging.⁵⁷

4.1 Surface grafting of polymers

The addition of sulfur-based (SB) functional groups leads to a decrease in hydrophobicity and roughness of the surface. Alves et al. grafted a polyurethane film with a sulfonic group, and the results of surface characterization tests and blood compatibility studies indicated an enhancement of the modified polyurethane biological performance, with increased blood compatibility.⁵⁸ Feng et al. tailored a polycarbonateurethane (PCU) surface with poly(ethylene glycol) monoacrylates (PEGMAs) with a molecular weight of between 400 and 1000 g mol⁽⁻¹⁾ to improve the hydrophilicity and haemocompatibility of the surface of the polycarbonateurethane (PCU). The surface-grafted PCU films were characterized by Fourier transformation infrared spectroscopy, X-ray photoelectron spectroscopy, contact angle, SEM, and atomic force microscopy measurements, and the blood compatibility of the surface was evaluated by platelet adhesion tests. The results showed that the hydrophilicity of the modified film had been improved significantly by grafting PEGMAs, and also that platelets adhesion onto the film surface was noticeably reduced. In addition, the molecular weight of the PEGMAs had a great influence on the hydrophilicity and haemocompatibility of the PCU films after surface modification, and increased with the increasing molecular weight of the PEGMAs.⁵⁹ Three zwitterionic polymers were grafted from a silicone rubber (SR) membrane. Observing the experimental results, all the zwitterionic polymer-modified surfaces have better resistance to protein adsorption and excellent resistance to platelet adhesion, thus showing significantly improved blood compatibility.⁶⁰ Poly(vinyl alcohol) (PVA) was added on a chitosan (CS) membrane surface, and biocompatibility was evaluated by FTIR, XRD and SEM examinations. The results suggest that adding PVA into a CS membrane could greatly improve the CS membrane's flexibility and wettability.⁶¹ Acetylated 1-thio-β-D-glucopyranose and 1-thio-β-D-galactopyranose were grafted onto a homopolymer of pentafluorostyrene (PFS) and onto a block copolymer of styrene and PFS. Finally the results showed that the grafted PFS are biocompatible for the 3T3 fibroblast and MC3T3-E1 preosteoblast cell lines.⁶² PEG cellulose is obtained by grafting PEG chains onto the cellulosic polymer. The results indicate that modified cellulose is an useful approach to improving the biocompatibility of the dialysis membrane for haemodialysis.⁶³ Sulfonated poly(ethylene oxide) (PEO) grafted polyurethane (PU) (PU-PEO-SO(3)) was examined using scanning electron microscopy and platelet adhesion; thrombus formation appeared to be appreciably lesser formed on the PU-PEO-SO(3) coated implants compared with the control PUs. The effectiveness of PU-PEO-SO(3)-coated implants in terms of blood compatibility, bio-stability and calcification resistance may make them a promising biomedical material in application for blood/tissue contacting implants and artificial organs.⁶⁴ The surface grafting of polymers has the advantage that the addition of a number of polymer chains on the polymer surface can be carried out easily, but the surface modification occurs through a reversible physical adsorption process, which is a drawback of the grafting technique.⁶⁵

4.2 Surface grafting of metals

The surface of stainless steel was modified by carbohydrate polymer grafting followed by acid-treatment. The surface investigation confirmed that the surface was changed from hydrophobic to hydrophilic and from rough to smooth. The biological experiments revealed that the surface-modified stainless steel not only inhibited non-specific fibrinogen adsorption, but also repelled most of the proteins from human blood. The treated stainless steel surfaces have improved biocompatibility when compared to bare stainless steel-based medical devices.⁶⁶ Kyomoto et al. grafted a 2-methacryloyloxyethyl phosphorylcholine (MPC) polymer onto the surface of a cobalt–chromium–molybdenum (Co–Cr–Mo) alloy to develop a highly biocompatible hip joint for total hip arthroplasty, and it was confirmed that the grafted metal surface had better biocompatibility than the raw cobalt alloy surface.⁶⁷ The alkaline phosphatase (ALP) enzyme was grafted to titanium and its alloy surface, and it was proved that the grafted metal surface showed improved bioactivity, as well as biocompatibility.⁶⁸ 1H,1H,2H,2H-Perfluorodecyl acrylate was added to the surface of diamond-like carbon (DLC)-deposited titanium metal, and the resulted surface was analyzed using X-ray photoelectron spectroscopy (XPS), contact angle measurement (CA), and a 3D surface profiler. All the results suggest that the biocompatibility and functional properties of the modified Ti₆Al₄V substrates were improved.⁶⁹ The addition of polymeric chains over the metal surface was a little complicated and the grafting was not durable, but it helped to attain the desired properties, such as corrosion and abrasion resistance.⁶⁷

4.3 Surface grafting of nanocomposites

The surfaces of BaTiO₃ nanoparticles were grafted with 1-tetradecylphosphonic acid (TDPA) to functionalize the surfaces. The acid-grafted nanoparticle surfaces were analyzed by FTIR, XPS and XRD. These results illustrated that the modified surfaces had improved flexibility and biocompatibility.⁷⁰ Cheng et al. modified the surface of a nano-hydroxyapatite fibrous scaffold with polyethylene glycol (PEG) to enhance the hydrophilicity of n-HA particles. The results of his study proved the improved wettability of the modified surface.⁷¹ The poly(ε-caprolactone) PCL-grafted HAp in nanocomposites provided a more favorable environment for protein adsorption and with better biocompatibility compared to unmodified HAp. Nanocomposites containing PCL-grafted nanophase HAp showed positive effects on fibroblast cell adhesion.⁷² The surfaces of BG nanoparticles were grafted with L-lactide to yield poly(L-lactide) (PLLA)-grafted gel particle (PLLA-g-BG). The nanocomposite with 20% PLLA-g-BG exhibited superior surface properties, including higher roughness and enhanced cell adhesion. The results show that the application of PLLA-g-BG with a certain blend ratio can improve the bioactivity and biocompatibility.⁷³ Iron/carbon nanoparticles (Fe@CNPs) are nanomaterials that are grafted with polymers, and the relationship between their biocompatibility and surface chemistry was investigated. The outcome of the investigation proved that the surface chemistry had a major effect on the biocompatibility of the grafted Fe@CNPs.⁷⁴ The surface grafting of nanocomposites avoids the aggregation of nanoparticles and can be used to form a stable suspension in organic solvents, but there are possibilities for negative effects on the surface chemistry.

5 Abrasive blasting and acid etching

Abrasive blasting is the process of forcibly propelling a stream of abrasive material against a surface under high pressure to smooth a rough surface, roughen a smooth surface, shape a surface, or to remove surface contaminants or some other substances from the material surface.⁷⁵ Grit blasting involves the projection of ceramic particles such as alumina, titanium oxide and calcium phosphate particles through a nozzle at high velocity by means of compressed air. Depending on the size of the ceramic particles, different surface roughnesses can be produced on medical implants. The blasting material should be chemically stable, biocompatible and should not produce negative effects on the material surface under treatment.^76,77 The clinical benefits in haemodialysis therapy is the removal of substances such as beta2-microglobulin (beta2-m), which has been reported by several authors. Additionally, the elimination of large-molecular weight “uremic toxins” is now generally acknowledged as being advantageous to the overall quality of life of patients by improving the membrane compatibility with human blood.⁷⁸ Kim projected the ceramic particles towards the surface of titanium implants at high velocity to obtain a high surface roughness. Then, the blasted surface of implants was modified by micro-arc oxidation treatment. A porous TiO₂ layer was formed on the surface and that could be attributed to the excellent biocompatibility.⁷⁹ The stable oxide layer over the metal surface plays a pivotal role in biocompatibility, and so an oxide layer was formed on the surface of stainless steel through grit-blasting, followed by micro-arc oxidation. The modified metal surface showed enhanced biocompatibility compared to the control group, and the modified stainless steel implant was suitable for cementless arthroplasty because of its outstanding biocompatibility due to the oxide layer formation.⁸⁰ Lampin et al. sandblasted the poly(methyl methacrylate) (PMMA) with alumina particles, and the treated surface was characterized in comparison with untreated samples. The results show that the sandblasted PMMA had an increased surface roughness, as well as showing the hydrophilicity of the polymer surface.⁸¹ Hossein et al. sandblasted the titanium (Ti₁₃Zr₁₃Nb) surface with alumina particles, followed by H₂SO₄ etching at 25 °C for 20 seconds. The SLA-treated surface was characterized with the aid of a field emission scanning electron microscope (FESEM), and the chemical composition was measured through energy dispersive X-ray (EDX) spectroscopy. The results of the FESEM and EDX show that the SLA-treated surface had better compatibility.⁸² Li et al. established a comparison study on various surface treatments of a titanium implant, such as sandblasting and acid etching, and finally ending with UV radiation. As a whole, UV irradiation was recognized as a trustworthy method for surface cleaning without changing the topography and roughness and led to a greater biocompatibility of the sandblasted and acid-etched titanium surface.⁸³ The titanium metal implant was treated with hydrofluoric acid solution (HF), and then the study of the modified surface displayed the increased roughness, lower cytotoxicity levels and better biocompatibility than the untreated implant surface.⁸⁴ The abrasive blasting technology involved in the removal of contaminants or other substances by forcibly propelling abrasive material may affect the mechanical properties of the material under modification, but it was the best choice for the surface cleaning of all types of biomaterials.^75,82

6 Surface coating

Coating is an effective method for surface modification to improve the biocompatibility of medical implants.⁸⁵ Various methods have been developed to coat medical materials, such as plasma spraying,^86,87 sputter deposition,^88,89 sol–gel coating,⁹⁰ electrophoretic deposition^91,92 biomimetic precipitation⁹³ and the dipping method;⁹⁴ the advantages and disadvantages of these coating methods are given in Table 1.

Table 1 List of merits and demerits of surface coating technologies

Coating technique	Thickness of coating layer	Merits	Demerits
Plasma spraying	50–250 μm	Can coat complex materials	Needs extremely high temperature, de-bonding of coated layer
Sputter deposition	0.02–1 μm	Uniform coating thickness	Expensive, time consuming, cannot coat complex surfaces
Sol–gel coating	<1 μm	Can coat complex shapes, low processing temperatures	Requires controlled atmosphere processing
Electrophoretic deposition	0.1–2.0 mm	Uniform coating thickness, rapid deposition rates, can coat complex materials	Difficult to produce crack-free coatings, requires high sintering temperatures
Biomimetic deposition	<30 μm	Coating of complex geometries, co-deposition of bio-molecules	Time consuming, requires controlled pH
Dipping method	0.05–0.5 mm	Inexpensive, coatings applied quickly, can coat complex substrates	Requires high sintering temperatures, thermal expansion mismatch

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Organophosphonic acids and organophosphonates are initially used for metal and metal oxide coatings for surface modification and for the modification of metal nanoparticles because of their inherent biocompatibility.⁹⁵ Mostly medical implants are coated (by plasma spraying or other methodologies) with layers of hydroxyapatite (HA), because it is more rapidly integrated into the human body than the other coating materials⁹⁶ and calcium phosphate to improve biocompatibility,⁹⁷ or mixtures of the two.⁹⁸ In the Bicon Implant System (Boston, USA), Star lock implants (Park Dental Research Corp, USA) and Osstem (South Korea), the surfaces are HA coated.⁹⁹ Since the application of plasma-sprayed hydroxyapatite coatings onto metallic bone implants in the 1980s, the concept of implant coatings has shifted from passively protecting thin films to active and instructive immobilized layers. Currently, a plethora of coating techniques is being investigated to actively coordinate a desired biological response at the interface between the artificial implants and the surrounding living tissue.¹⁰⁰ Butruk et al. modified the polyurethane (PU) surface with soybean-derived phosphatidylcholine (PC) by a one-step dip-coating technique. To estimate the blood compatibility of the resulting material, modified samples were contacted with human blood. The PC-treated surfaces were thoroughly analyzed and tested for fibrinogen resistance, the ability to oppose platelet adhesion, haemolysis ratio and plasma re-calcification time. The outcomes of this analysis demonstrated a significant reduction in fibrinogen deposition to PC-modified materials, as compared to non-modified PU. The proportion of non-aggregated platelets remaining in the blood samples contacted with PC-coated materials exceeded 70%. The same parameter measured for the control PU was significantly lower, and was about 28%.⁹³ The medical polymer polyurethane was coated with polyaniline (PANi) and a polyaniline–silver nanoparticle composite (PANi–AgNp) and the coated surface characteristics were investigated. Contact angle measurements indicated that the hydrophilic surfaces are compatible to cells when compared to unmodified surfaces. These modifications make the surface more biocompatible than the original PU. The coating of polymers is easy to implement, but it decreases the resistance to heat and the coating layer is not durable.¹⁰¹ HA-coated bone implants have improved biological fixation, and show better fixation after 4 weeks of implantation. It can be concluded that the HA coating is an effective method for improving bone formation for orthopaedic implants with enhanced biocompatibility.¹⁰² A titanium oxide ceramic coating of 2000 Å to 2500 Å thickness on the titanium implant surface was subjected to heat treatment to enhance the ceramic coating adherence with the metal surface. The resulting coated metal surface showed higher surface hardness and was suitable for orthopaedic and dental implants.¹⁰³ The ceramic coating increases the strength of the chemical bonding holding the atoms and molecules together and thus further improves the hardness of the material.¹⁰⁴ The gelatin nano gold (GnG) composite is used for the surface coating of titanium in addition to ensure biocompatibility. Surface characterization tests were performed to evaluate the haemocompatibility of the modified surface and the results show that the GnG-coated surface had better compatibility than the pure titanium.¹⁰⁵ Kim et al. coated the magnesium (Mg) surface with hydroxyapatite (HA) in an aqueous solution containing calcium and phosphate sources to improve its in vitro and in vivo bio-corrosion resistance, biocompatibility and bone response. The preliminary in vivo experiments also showed that the bio-corrosion of the Mg implant was significantly retarded by the HA coating, which results in good mechanical stability and improved biocompatibility.¹⁰⁶ A surface coating of poly(1,3-trimethylene carbonate) (PTMC) on a magnesium (Mg) alloy was investigated. The haemocompatibility and histocompatibility of the coated surface were examined and compared with a control sample. The results revealed that the PTMC-coated surface led to less haemolysis than on the controls.¹⁰⁷ The nanocomposites of fibronectin (FN) and gold nanoparticles AuNPs (FN–Au) were surface modified and analyzed by atomic force microscopy (AFM), UV-Vis spectrophotometry (UV-Vis), and Fourier transform infrared spectroscopy (FTIR). The biocompatibility of the nanocomposites was assessed by the response of monocytes and platelets to the material surface in vitro. These results suggest that the FN–Au nanocomposite thin film coating may serve as a potential and uncomplicated solution for the surface modification of blood-contacting medical devices.¹⁰⁸ For all types of raw materials, the advantages of the coating technique include the deposition of holes present in the surface and a low processing temperature, but the selection of the coating material still remains a challenge.^95,104

7 Thermal treatment

Heat treating is a group of industrial and metal working processes used to alter the physical and chemical properties of a material. Kawase et al. made a comparison study on control platelet-rich fibrin (PRF) with heat-compressed PRF. The heat-treated sample appeared plasmin-resistant and remained stable for longer than 10 days. Moreover, in animal implantation studies, the heat-compressed PRF was experimented on for at least three weeks after implantation in vivo; whereas, the control PRF was completely degraded within two weeks. Therefore, these findings suggest that the heat-compression technique decreases the rate of biodegradation of the PRF membrane, without sacrificing its biocompatibility.¹⁰⁹ Titanium was modified by means of a hydrothermal treatment with a maximum pressure of 6.3 MPa at 280 °C in calcium oxide (CaO) solution or water to improve the bioactivity and biocompatibility. As a result, calcium titanate was formed on the titanium surface, which showed improved bioactivity and biocompatibility.¹¹⁰ Titanium was also treated with NaOH solution at 60 °C for 24 hours, followed by heat treatment up to 600 °C for 1 h. The results infer that the alkali solution heat-treated surface had improved biocompatibility.¹¹¹ The titanium implant was sintered with tricalcium phosphate (TCP) by spark plasma at 1200 °C; the final TCP–Ti composite showed improved cell viability and proliferation. An in vivo study confirmed that within three months of implantation in an animal body, 70% TCP–Ti had an excellent bone-implant interface compared with a pure Ti metal implant.¹¹² In addition to bioactivity, orthopaedic implants require porosity for tissue regeneration; heating at high temperatures (500–1000 °C) resulted in porosity and directly led to positive consequent modifications in the mechanical properties and biocompatibility, biodegradability and bioactivity of the material surface.^113,114 The effects of heat treatment on the alloys were studied and it was documented that the microstructure of an alloy was changed due to the thermal effect; the end results could be useful in further understanding the relationship between the biocompatibility, wear and corrosion resistance of the alloy, so as to allow the development of a promising biomedical material.¹¹⁵ Bimbo et al. demonstrated that thermally hydrocarbonized porous silicon (THCPSi) nanoparticles did not induce any significant toxicity, oxidative stress, or inflammatory response in Caco-2 and RAW 264.7 macrophage cells. On the whole, these silicon-based nanosystems exhibit outstanding in vivo stability, biocompatibility, low cytotoxicity and non-immunogenic profiles, making them ideal for oral drug delivery purposes.¹¹⁶ Cui et al. applied hot water and heat treatments to transform the titania layers from an amorphous structure into a crystalline structure with enhanced compatibility. The loads of Ti–OH groups formed by hot water treatment could contribute to apatite formation on the surface of titanium metals, and subsequent heat treatment would enhance the bond strength between the apatite layers and the titanium substrates. Thus, bioactive titanium metals could be prepared via hot water, followed by a heat treatment that would be suitable for applications under load-bearing conditions.¹¹⁷ However, the high operating temperature of the thermal modification is a major drawback. The polymeric and ceramic materials failed to withstand this high level temperature and so thermal treatment is mostly preferred for metallic biomaterials possessing a high melting point that leads to optimistic changes in the mechanical and physico-chemical properties of metals.

8 Vapour and steam treatment

Water vapour is little water droplets that exist in the air, while steam is water heated to the point that it turns into gas. In simplified science, both are referred to as the gaseous state of water. Steam is believed to be basically water vapour at a higher temperature. A vapour is a matter in the gas state at a temperature below its critical point, and it is used in the field of biomaterial science to improve the blood compatibility of biomaterials.¹¹⁸ Jensen et al. etched polydimethylsiloxane (PDMS) for 8 min in water vapour at a pressure of 50 m Torr and power of 400 W, which resulted in unwavering long-term wettability and excellent in vitro cell compatibility. Ultimately, it was concluded that water vapour plasma may be used to improve the biointegration of PDMS implants and thereby evade clinical problems related with the formation of a dead space.¹¹⁹ Silicon carbide (SiC), a chemical vapour deposition coating for cardiovascular implants, resulted in a decline in platelet adhesion and also less inflammatory reactions. Diamond-like carbon has comparable advantages as SiC and also offers higher hardness, lower frictional coefficient, chemical inertness, bio-stability, and also good blood compatibility, making it a graceful alternative for the application on vascular stents.¹²⁰ Wang et al. used human haepatoma cells (BEL-7402) as model cells to examine cell adhesion, spreading and the proliferation of cells on zein films before and after surface treatment with water vapour, and he concluded that hydrophilicity and cell adhesion were significantly improved after the treatment on zein films.¹²¹ Non-woven polyethylene terephthalate (NW-PET) was subjected to surface modification under water vapour to enhance the compatibility, and the outcome of this study illustrated that the water-vapour-treated NW-PET had improved platelet compatibility.¹²² The hydroxyapatite (HA) coatings were kept in water vapour at 125 °C, with a pressure of 0.15 MPa for 6 h to modify the amorphous phase in the coating into crystalline HA and this improved the stability of the coating.¹²³ Lee et al. established that the water vapour treatment is an easy and valuable technique to fabricate hydroxyl groups on polymer surfaces, which possibly has a positive effect on cell adherence with increased wettability.¹²⁴

Steam is the technical term for the gaseous phase of water, which is formed when water boils, and is a frequently used surface treatment for biomaterials to seal the interconnected pores and to enhance the superficial properties of the material. The treatment with steam forms an oxide layer on the material surface that fills the pores on the surface, thereby increasing the density, and also has an effect on the hardness of the material.¹²⁵ Steam treating is the controlled oxidation of metals to produce a thin layer of oxide on the surface of a component. This process can be used to provide a component with improved corrosion resistance, better wear resistance, increased surface hardness, and wettability.¹²⁶ The ferrous components were subjected to steam treatment to improve their mechanical properties through the formation of an oxide layer, which gives better biocompatibility; also, the reduced heat required for steam production and its environmental benefits makes steam treatment technology a growing application in surface modification.¹²⁷ The high biocompatibility of Ti and the alloys of Ti is due to the formation of an oxide layer during the process of implant preparation and can be modified by steam sterilization, resulting in increased oxide layer thickness with respect to the unsterilized samples.¹²⁸ Rai et al. sterilized the poly(glycerol sebacate) (PGS) with steam and concluded that the sterilized samples maintain their mechanical properties and compatibility, and the treated PGS was used for wider applications in medical devices.¹²⁹ Chen et al. performed a hydrophilization treatment on a graphene surface using steam to reduce the interfacial impedance during cardiac and neural recording through converting it into a hydrophilic nature, which results in enhanced biocompatibility.¹³⁰ The mechanical properties of chitosan/hydroxyapatite (HAp) nanocomposites were improved by heat treatment with saturated steam, which led to hydrogen bond formation between the chitosan molecules. The treated nanocomposites were implanted into rats and after three weeks of post implantation, it was found that the cells were seen around the composite accompanied with a surface roughness showing enhanced biocompatibility.¹³¹ The limitations of vapour and steam treatment include that the material under the surface modification should tolerate the heat of the vapour or steam. However, it is easy, cost effective and there is no need for the further sterilization of modified surface, and thus these are the merits of vapour and steam treatment.¹²⁴

9 Conclusion

The surface properties and biocompatibility of the biomaterials were improved by various surface modification techniques, namely, physico-chemical, mechanical and biological methods. Among these three techniques, the physico-chemical method has been used as a versatile strategy in recent days. This method includes both physical and chemical means of treatments, mostly the addition of some reagents to the surface and heat treatment. The merits and demerits of each physico-chemical modification technology for each raw material are listed in Table 2. From the results and the remarks from various works, we can conclude that most research has been aimed at chemical modifications to improve the biocompatibility by altering the surface characteristics of the medical material.

Table 2 List of merits and demerits of surface modification technologies for each raw material

Modification technology	Type of material	Merits	Demerits
Chemical modification	Polymers	Produces new polymeric material which cannot be synthesized by a polymerization reaction	Non specific interaction
	Metals	Desired topographical properties	Effects based on chemical reagents or molecules used for the modification
	Ceramics	Increased bioactivity	Selection of chemical reagent remains challenging
	Nanocomposites	Reduction in agglomeration	Effects based on chemical reagents or molecules used for the modification
Surface grafting	Polymers	Controllable addition of polymeric chains	Physical adsorption taking place is reversible
	Metals	Better corrosion resistance	Non specific localization of grafted molecules
		Improved abrasion resistance	Complicated process
	Nanocomposites	Aggregation of nanoparticles are avoided	Possibilities of negative effects on the surface chemistry
		Forms a stable suspension in organic solvents
Abrasive blasting and acid etching	Polymers, metals, ceramics and nanocomposites	Removes contaminants	Forcibly propelling of abrasive material may damage the surface
		Best choice of surface cleaning	Mechanical properties may be affected
Surface coating	Polymers	Deposition of holes	Decreased heat resistance
		Easy to implement
	Metals	Deposition of holes	Coating layer is not durable
		Better corrosion resistance
	Ceramics	Increases the strength of chemical bonding	Selection of coating material is complicated
		Increased hardness
	Nanocomposites	Enhanced interaction of coated material with the surface	Non stable coating layer
Thermal treatment	Polymers, metals, ceramics and nanocomposites	Optimistic changes in mechanical properties	Very high operating temperature
		Improved physico-chemical properties	Only suitable for materials with a high melting point
Vapour and steam treatment	Polymers, metals, ceramics and nanocomposites	Cost effective	Production of high temperature
		No need for further sterilization	Produce adverse thermal effects

---

Mostly all modalities under physico-chemical modification were performed on metals and their alloys, due to the high mechanical strength of metallic bonding that can withstand high temperature without losing its shape. The surface treatment of polymers can be mainly achieved by surface grafting and coating mechanisms. However, further experimentation of polymers using other modalities may further promote it as a promising biomaterial for various biomedical applications. The recent generation of biomaterials, such as ceramics and nanocomposites, has not yet been subjected to much surface modification because of their tailor-made characteristics for particular biomedical applications. Despite its pre-defined characteristics, there are some deficits with biocompatibility, so ceramics and nanocomposites are still expected to be investigated in depth in order to gain insights about the prospect of physico-chemical treatments in these materials.

Physico-chemical modifications induce changes in the physical properties of all biomaterials, like modification in their surface roughness and wettability. These physical changes due to the physico-chemical modification are graphically represented in Fig. 2. These physical variations lead to improved biocompatibility, decreased platelet adhesion, enhanced protein adsorption and reduced red blood cell damage, as shown in Fig. 3. Biomedical materials subjected to physico-chemical modification are a more attractive choice for diverse applications like cardiovascular, tissue regeneration and orthopaedic applications, since the surfaces can be optimized for the particular applications. Hence, proper exploitation of this strategy will quench the thirst of long-time unmet demands for biocompatibility.

Fig. 2 Physical changes of the biomaterial surface due to physico-chemical modification.

Fig. 3 Enhanced biocompatibility changes induced by the physico-chemical modification.

Acknowledgements

This work was supported partly by the Ministry of Higher Education Malaysia with the Grant Vot no: R.J130000.7809.4F444 and also acknowledges the support of UPMU, UTM.

Notes and references

1. D. Bronzino, The Biomedical Engineering Handbook, 2nd edn, 2000, vol. 1.
2. http://reportlinker.com/ci02234/Biomaterial.html.
3. P. Parida, A. Behera and S. Chandra Mishra, Int. J. Adv. Appl. Sci., 2012, 125–129.
4. B. Singh and L. Pal, J. Mech. Behav. Biomed. Mater., 2012, 9–21.
5. G. Manivasagam, D. Dhinasekaran and A. Rajamanickam, Recent Pat. Corros. Sci., 2010, 40–54.
6. D. Williams, Medical Device Technologies, 2003, 14(8), 10–13.
7. J. Lemmons and J. Natiella, Dent. Clin. North Am., 1986, 30, 3–23.
8. G. Schmalz, J. Dent. Res., 2002, 81, 660–663.
9. S. K. Jaganathan, E. Supriyanto, S. Murugesan, A. Balaji and M. K. Asokan, BioMed. Res. Int., 2014, 1–11.
10. M. Shuchman, N. Engl. J. Med., 2006, 355, 1949–1952.
11. D. F. Williams, Biomaterials, 2008, 29, 2941–2953.
12. E. Gentleman, M. D. Ball and M. M. Stevens, Medical Sciences – Vol II – Biomaterials, EOLSS publishers, Oxford, UK, 2009.
13. S. K. Jaganathan, A. Balaji, M. V. Vellayappan, A. P. Subramanian, A. A. John, M. K. Asokan and E. Supriyanto, J. Mater. Sci., 2014, 2007–2018.
14. R. L. McCreery and A. J. Bergren, in Nanofabrication, ed. M. Stepanova and S. Dew, Springer-Verlag, Vienna, 2012, pp. 163–190,  DOI:10.1007/978-3-7091-0424-8_7.
15. V. Safonov, A. Zykova, J. Smolik, R. Rogovska, N. Donkov and V. Georgieva, J. Phys.: Conf. Ser., 2010 DOI:10.1088/1742-6596/253/1/012068.
16. I. Milinkovi, R. Rudolf, K. T. Rai, Z. Aleksi, V. Lazi, A. Todorovi and D. Stamenkovi, Mater. Technol., 2012, 251–256.
17. F. Variola, J. B. Brunski, G. Orsini, P. Tambasco de Oliveira, R. Wazenm and A. Nanci, Nanoscale, 2011, 3, 335–353.
18. F. Abbasi, H. Mirzadeh and A.-A. Katbab, Polym. Int., 2002, 51, 882–888.
19. K. N. J. Stevens, PhD thesis, Maastricht University, 2011.
20. A. P. Alekhin, G. M. Boleiko, S. A. Gudkova, A. M. Markeev, A. A. Sigarev, V. F. Toknova, A. G. Kirilenko, R. V. Lapshin, E. N. Kozlov and D. V. Tetyukhin, Nanotechnol. Russ., 2010, 5, 696–708.
21. S. Bertazzo, W. F. Zambuzzi, H. A. da Silva, C. V. Ferreira and C. A. Bertran, Clin. Oral Implants Res., 2009, 20, 288–293.
22. L. Gabo, K.-Y. Chen, G. T. Carroll and B. L. Feringa, Chem.–Eur. J., 2013, 19(32), 10690–10697.
23. A. S. Hoffman, Chin. J. Polym. Sci., 1995, 3, 13.
24. J. Y. Lim and H. J. Donahue, Tissue Eng., 2007, 13(8), 1879–1891.
25. A. Thapa, D. C. Miller, T. J. Webster and K. M. Haberstroh, Biomaterials, 2003, 24(17), 2915–2926.
26. D. C. Miller, A. Thapa, K. M. Haberstroh and T. J. Webster, Biomaterials, 2004, 25(1), 53–61.
27. B. Balakrishnan, D. S. Kumar, Y. Yoshida and A. Jayakrishnan, Biomaterials, 2005, 26(17), 3495–3502.
28. Y. B. J. Aldenhoff and L. H. Koole, Eur. Cells Mater., 2003, 5, 61–67.
29. T. Xiang, W. W. Yue, R. Wang, S. Liang, S. D. Sun and C. S. Zhao, Colloids Surf., B, 2013, 110, 15–21.
30. S. Nie, J. Xue, Y. Lu, Y. Liu, D. Wang, S. Sun, F. Ran and C. Zhao, Colloids Surf., B, 2012, 116–125.
31. S. K. Jaganathan, H. Mohandas, G. Sivakumar, P. Kasi, T. Sudheer, S. Avineri Veetil, S. Murugesan and E. Supriyanto, BioMed. Res. Int., 2014, 1–7.
32. H. D. Park, W. K. Lee, T. Ooya, K. D. Park, Y. H. Kim and N. Yui, J. Biomed. Mater. Res., Part A, 2003, 66(3), 596–604.
33. Z. K. Xu, F. Q. Nie, C. Qu, L. S. Wan, J. Wu and K. Yao, Biomaterials, 2005, 26(6), 589–598.
34. A. Akelah, Technology and Engineering, 2013, 1–355,  DOI:10.1007/978-1-4614-7061-8.
35. K. M. Hansson, S. Tosatti and J. Isaksson, et al., Biomaterials, 2005, 26, 861–872.
36. M. Danila Korotin, S. Bartkowski, E. Z. Kurmaev, M. Meumann, B. Eugeniya Yakushina, R. Z. Valiev and O. S. Cholakh, J. Biomater. Nanobiotechnol., 2012, 3, 87–91.
37. W. Xue, X. Liu, X. Zhen and C. Ding, Biomaterials, 2005, 26(16), 3029–3037.
38. X. Zhao, X. Liu, C. Ding and P. K. Chu, Surf. Coat. Technol., 2006, 5487–5492.
39. C. H. Yea, Y. F. Zhenga, S. Q. Wang, T. F. Xia and Y. D. Li, Appl. Surf. Sci., 2012, 258, 3420–3427.
40. M. H. Lee, D. Yoon, D. Won, T. Bae and F. Watari, Met. Mater. Int., 2003, 9, 35–42.
41. A. Parsapour, S. Nouri Khorasani and M. Hossein Fathi, J. Mater. Sci. Technol., 2012, 28(2), 125–131.
42. H. Zreiqat, P. Evans and C. R. Howlett, J. Biomed. Res., 1999, 44, 389–396.
43. K. Nishizawa, M. Toriyama, T. Suzuki, Y. Kawamoto, Y. Y. Okogawa and H. Nagae, J. Ferment. Bioeng., 1993, 75, 435–437.
44. Q. Zhao, G. J. Zhai, D. H. L. Ng, X. Z. Zhang and Z. Q. Chen, Biomaterials, 1999, 20, 595–599.
45. W. Zhang, W. Liu, W. Gu, L. Chen and Y. Shen, Adv. Mater. Res., 2012, 391–392, 195–199.
46. S. Hsua, H. Hsua, W. F. Hob, C. H. Yaoc, P. Changd and S. C. Wu, The 41st International Conference on Metallurgical Coatings and Thin Films, 2014, vol. 572, pp. 91–98.
47. B. Duan and M. Wang, J. R. Soc., Interface, 2010, 7, S615–S629.
48. C. Deng, X. Xiao, N. Yao, X. B. Yang and J. Weng, Int. J. Polym. Mater. Polym. Biomater., 2011, 60(12), 969–978.
49. M. C. Pascu, M. C. Popescu and C. Vasile, J. Phys. D: Appl. Phys., 2008, 41, 12.
50. M. Selvakumar, S. K. Jaganathan, G. B. Nando and S. Chattopadhyay, J. Biomed. Nanotechnol., 2015, 11, 291–305.
51. A. R. Adhikari, I. Rusakova, A. Haleh, J. Luisi, N. I. Panova, F. Laezza and W. K. Chu, J. Appl. Phys., 2014, 115.
52. Q. Yue, M. Wang, Z. Sun, C. Wang, Y. Deng and D. Zhao, J. Mater. Chem. B, 2013, 1, 6085–6093.
53. A. Mohammadi, M. Barikani and M. Barmar, J. Mater. Sci., 2013, 48, 7493–7502.
54. A. V. Jovanovic, J. A. Flint, M. Varshney, T. E. Morey, D. M. Dennis and R. S. Duran, Biomacromolecules, 2013, 7(3), 945–949.
55. X. H. Gong, C. Y. Tang, H. C. Hu, X. P. Zhou and X. L. Xie, J. Mater. Sci.: Mater. Med., 2004, 15(10), 1141–1146.
56. L. S. Wang and R. Y. Hong, Adv. Nanocompos., 2011, 289–322.
57. S. Minko, Polym. Surf. Interfaces, 2008, 215–234.
58. P. Alves, S. Pinto, P. Ferreira, J. P. Kaiser, A. Bruinink, H. C. de Sousa and M. H. Gil, J. Mater. Sci.: Mater. Med., 2014, 2017–2026.
59. Y. Feng, H. Zhao, M. Behl, A. Lendlein, J. Guo and D. Yang, J. Mater. Sci.: Mater. Med., 2013, 24(1), 61–70.
60. P. Liu, Q. Chen, B. Yuan, M. Chen, S. Wu, S. Lin and J. Shen, Mater. Sci. Eng., C, 2013, 33(7), 3865–3874.
61. P. Zhuanga, Y. Lia, L. Fanb, J. Linb and Q. Hu, Int. J. Biol. Macromol., 2012, 50, 658–663.
62. K. Babiuch, C. R. Becer, M. Gottschaldt, J. T. Delaney, J. Weisser, B. Beer, R. Wyrwa, M. Schnabelrauch and U. S. Schubert, Macromol. Biosci., 2011, 11, 535–548.
63. V. Sirolli, S. Di Stante, S. Stuard, L. Di Liberato, L. Amoroso, P. Cappelli and M. Bonomini, Int. J. Artif. Organs, 2000, 23(6), 356–364.
64. L. Yang, J. C. Davis, R. Vankayala, K. C. Hwang, J. Zhao and B. Yan, Biomaterials, 2010, 31, 5083–5090.
65. M. Ebara, et al., Smart Biomaterials, 2014.
66. C. Kanga and Y. Lee, J. Ind. Eng. Chem., 2012, 18, 1670–1675.
67. M. Kyomoto, T. Moro, Y. Iwasaki, F. Miyaji, H. Kawaguchi, Y. Takatori, K. Nakamura and K. Ishihara, J. Biomed. Mater. Res., Part A, 2009, 91, 730–741.
68. S. Ferraris, S. Spriano, C. L. Bianchi, C. Cassinelli and E. Vernè, J. Mater. Sci.: Mater. Med., 2011, 22, 1835–1842.
69. H. D. Park, W. K. Lee, T. Ooya, K. D. Park, Y. H. Kim and N. Yui, J. Biomed. Mater. Res., Part A, 2003, 66(3), 596–604.
70. H. J. Ye, W. Z. Shao and L. Zhen, Colloids Surf., A, 2013, 427, 19–25.
71. Z. Cheng, G. Pang, H. Wang, J. Li and X. Zhao, Adv. Mater. Res., 2012, 1095–1099.
72. H. J. Lee, S. E. Kim, H. W. Choi, C. W. Kim, K. J. Kim and S. C. Lee, Eur. Polym. J., 2007, 43, 1602–1608.
73. S. J. Dong, T. Yu, J. C. Wei, X. B. Jing, Y. M. Zhou, P. B. Zhang and X. S. Chen, Chem. J. Chin., 2009, 30, 1018–1023.
74. Y. Liu, X. Fu, Y. Bu, J. Zhang, J. Zhang and W. Longmin, Surf. Coat. Technol., 2012, 208, 51–56.
75. B. Pattanaik, S. Pawar and S. Pattanaik, Indian Journal of Dental Research, 2012, 23(3), 398–406.
76. A. Wennerberg, T. Albrektsson and B. Andersson, Int. J. Oral Maxillofac. Implants, 1996, 11, 38–45.
77. A. Citeau, J. Guicheux, C. Vinatier, P. Layrolle, T. P. Nguyen and P. Pilet, et al., Biomaterials, 2005, 26, 157–165.
78. S. K. Bowry, Int. J. Artif. Organs, 2002, 25(5), 447–460.
79. Y. W. Kim, Mater. Manuf. Processes, 2010, 25, 307–310.
80. Y. M. Lim, S. Y. Kwon, D. H. Sun and Y. S. Kim, Clin. Orthop. Relat. Res., 2011, 469, 330–338.
81. M. Lampin, R. Warocquier-Clérout, C. Legris, M. Degrange and M. F. Sigot-Luizard, J. Biomed. Mater. Res., 1997, 36, 99–108.
82. H. M. Khanlou, Afr. J. Basic Appl. Sci., 2012, 6(6), 125–131.
83. S. Li, J. Ni, X. Liu, X. Zhang, S. Yin, M. Rong, Z. Guo and L. Zhou, J. Biomed. Mater. Res., Part B, 2012, 100, 1587–1598.
84. S. F. Lamolle, M. Monjo, M. Rubert, H. J. Haugen, S. P. Lyngstadaas and J. E. Ellingsen, Biomaterials, 2009,(5), 736–742.
85. T. Shirzadian, S. Bagheri, H. SaeidiBorojeni, P. Ghaffari, F. Foroughi and M. Mahboubi, J. Inj. Violence Res., 2012, 4(3), 39–40.
86. Z. Strnad, J. Strnad, C. Povýsil and K. Urban, Int. J. Oral Maxillofac. Implants, 2000, 15, 483–490.
87. M. Gottlander, C. B. Johansson and T. Albrektsson, Clin. Oral Implants Res., 1997, 8, 345–355.
88. M. Yoshinari, Y. Ohtsuka and T. Dérand, Biomaterials, 1994, 15, 529–535.
89. Y. Yang, K. H. Kim and J. L. Ong, Biomaterials, 2005, 26, 327–337.
90. H. Q. Nguyen, D. A. Deporter, R. M. Pilliar, N. Valiquette and R. Yakubovich, Biomaterials, 2004, 25, 865–876.
91. D. Lakstein, W. Kopelovitch, Z. Barkay, M. Bahaa, D. Hendel and N. Eliaz, Acta Biomater., 2009, 5, 2258–2269.
92. J. Wang, P. Layrolle, M. Stigter and K. de Groot, Biomaterials, 2004, 25, 583–592.
93. F. Barrère, P. Layrolle, C. A. Van Blitterswijk and K. De Groot, J. Mater. Sci. Mater. Med., 2001, 12, 529–534.
94. B. Butruk-Raszeja, M. Trzaskowski and T. Ciach, J. Biomater. Appl., 2014, 801–812.
95. C. Queffélec, M. Petit, P. Janvier, D. Andrew Knight and B. Bujoli, Chem. Rev., 2012, 112, 3777–3807.
96. M. Niinomi, M. Nakai, J. Hieda, K. Cho, T. Kasuga, T. Hattori, T. Goto and T. Hanawa, Int. J. Surf. Sci. Eng., 2012, 8, 138–152.
97. K. De Groot, J. Biomed. Mater. Res., 1989, 23, 1367–1371.
98. H. W. Kim, G. Georgiou, J. C. Knowles, Y. H. Koh and H. E. Kim, Biomaterials, 2004, 25(18), 4203–4213.
99. E. J. Lee, S. H. Lee, H. W. Kim, Y. M. Kong and H. E. Kim, Biomaterials, 2005, 26, 3843–3851.
100. R. Bosco, J. Beucken, S. Leeuwenburgh and J. Jansen, Coatings, 2012, 2, 95–119.
101. P. K. Prabhakar, S. Raj, P. R. Anuradha, S. N. Sawant and M. Doble, Colloids Surf., B, 2011, 86, 146–153.
102. A. Moroni, V. L. Caja, E. L. Egger, L. Trinchese and E. Y. Chao, Biomaterials, 1994, 15(11), 926–930.
103. G. Szabó, L. Kovács, K. Vargha, J. Barabás and N. Z. Émeth, J. Long-Term Eff. Med. Implants, 1999, 9, 247–259.
104. I. Milinkovi, R. Rudolf, K. T. Rai, Z. Aleks, V. Lazi, A. Todorovi and D. Stamenkovi, Mater. Technol., 2012, 3, 251–256.
105. Y. H. Lee, G. Bhattaraia and S. Aryal, et al., Appl. Surf. Sci., 2010, 256, 5882–5887.
106. S. M. Kim, J. H. Jo, S. M. Lee, M. H. Kang, H. E. Kim, Y. Estrin, J. H. Lee, J. W. Lee and Y. H. Koh, J. Biomed. Mater. Res., Part A, 2014, 102(2), 429–441.
107. J. Wanga, Y. Hea and M. F. Maitz, et al., Acta Biomater., 2013, 9, 8678–8689.
108. H. S. Hung, C. M. Tang, C. H. Lin, S. Z. Lin and M. Y. Chu, et al., PLoS One, 2013, 8, 65738,  DOI:10.1371/journal.pone.0065738.
109. T. Kawase, M. Kamiya, M. Kobayashi, T. Tanaka, K. Okuda, L. F. Wolff and H. Yoshie, J. Biomed. Mater. Res., Part B, 2014 DOI:10.1002/jbm.b.33262.
110. R. Sultana, M. Kon, L. M. Hirakata, E. Fujihara, K. Asaoka and T. Ichikawa, Dent. Mater. J., 2006, 25(3), 470–479.
111. K. Feng, E. Wu, Y. Pan and K. Ou, Mat Trans., 2007, 48, 2978–2985.
112. D. Mondal, L. Nguyen, I. H. Oh and B. T. Lee, J. Biomed. Mater. Res., Part A, 2013, 101(5), 1489–1501.
113. P. Majumdar, S. B. Singh, S. Dhara and M. Chakraborty, J. Mech. Behav. Biomed. Mater., 2012, 10, 1–12.
114. Y. Zhou, H. Li, K. Lin, W. Zhai, W. Gu and J. Chang, J. Mater. Sci.: Mater. Med., 2012, 23(9), 2101–2108.
115. C. Z. Wu, S. C. Chen, Y. H. Shih, J. M. Hung, C. C. Lin, L. H. Lin and K. L. Ou, J. Mech. Behav. Biomed. Mater., 2011, 4(7), 1548–1553.
116. L. M. Bimbo, M. Sarparanta, H. A. Santos, A. J. Airaksinen, E. Mäkilä, T. Laaksonen, L. Peltonen, V. P. Lehto, J. Hirvonen and J. Salonen, ACS Nano, 2010, 4(6), 3023–3032.
117. X. Cui, H. M. Kim, M. Kawashita, L. Wang, T. Xiong, T. Kokubo and T. Nakamura, J. Mater. Sci.: Mater. Med., 2008, 19(4), 1767–1773.
118. M. J. Jackson, G. M. Robinson, N. Ali, Y. Kousar, S. Mei, J. Gracio, H. Taylor and W. Ahmed, J. Med. Eng. Technol., 2006, 30(5), 323–329.
119. C. Jensen, L. Gurevich, A. Patriciu, J. J. Struijk, V. Zachar and C. P. Pennisi, J. Biomed. Mater. Res., Part A, 2012, 100, 3400–3407.
120. M. Fedel, A. Motta, D. Maniglio and C. Migliaresi, J. Biomed. Mater. Res., Part B, 2009, 90, 338–349.
121. H. J. Wang, J. X. Fu and J. Y. Wang, Colloids Surf., B, 2009, 69(1), 109–115.
122. E. H. Kostelijk, A. J. Klomp, G. H. Engbers, C. W. Gouwerok, A. J. Verhoeven, W. G. Van Aken, J. Feijen and D. de Korte, Prog. Transfus. Med., 2001, 11(3), 199–205.
123. Y. Cao, J. Weng, J. Chen, J. Feng, Z. Yang and X. Zhang, Biomaterials, 1996, 17(4), 419–424.
124. J. H. Lee, J. W. Park and H. B. Lee, Biomaterials, 1991, 12(5), 443–448.
125. P. Nader, H. Zahra and N. Mohammad, Comat Recent Trends in Structural Materials, 2012, pp. 1–6.
126. Y. Cao, F. Chan, Y. H. Chui and H. Xiao, BioResources, 2012, 7, 4109–4121.
127. S. L. Feldbauer, Steam Treating; Enhancing the surface properties of metal components.
128. N. L. Hernández de Gatica, G. L. Jones and J. A. Gardella Jr, Appl. Surf. Sci., 1993, 68, 107–121.
129. R. Rai, M. Tallawi, J. A. Roether, R. Detsch, N. Barbani, E. Rosellini, J. Kaschta, D. W. Schubert and A. R. Boccaccini, Mater. Lett., 2013, 105, 32–35.
130. C. H. Chen, C. T. Lin, W. L. Hsu, Y. C. Chang, S. R. Yeh, L. J. Li and D. J. Yao, Nanomedicine, 2013, 9(5), 600–604.
131. H. Kashiwazaki, Y. Kishiya, A. Matsuda, K. Yamaguchi, T. Iizuka, J. Tanaka and N. Inoue, Bio-Med. Mater. Eng., 2009, 19, 133–140.

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