J. Biomater. Sci. Polymer Edn, Vol. 17, No. 11, pp. 1241– 1268 (2006)  VSP 2006. Also available online - www.brill.nl/jbs Review Nanobiomaterials: a review of the existing science and technology, and new approaches V. HASIRCI 1,2,∗ , E. VRANA 1,2 , P. ZORLUTUNA 1,2 , A. NDREU 1,2 , P. YILGOR 1,2 , F. B. BASMANAV 1,2 and E. AYDIN 1,2 1 METU, 2 METU, Department of Biological Sciences, Biotechnology Research Unit, Ankara 06531, Turkey Department of Biotechnology, Biotechnology Research Unit, Ankara 06531, Turkey Received 30 January 2006; accepted 26 May 2006 Abstract—Nanotechnology has made great strides forward in the creation of new surfaces, new materials and new forms which also find application in the biomedical field. Traditional biomedical applications started benefiting from the use nanotechnology in an array of areas, such as biosensors, tissue engineering, controlled release systems, intelligent systems and nanocomposites used in implant design. In this manuscript a review of developments in these areas will be provided along with some applications from our laboratories. Key words: Nanobiomaterials; tissue engineering; drug delivery; composites; nanofibers. GUIDED TISSUE ENGINEERING AND NANOPATTERNING Cells respond to chemical and physical cues, and they show strong responses to the features of a biomaterial [1]. These may manifest themselves in various ways such as strong adhesion, detachment, cell spreading or migration. Chemical cues are mainly the presence of certain molecules that influence adhesion of cells such as proteins laminin and fibronectin, or those that influence the hydrophilicity of the material involved [2]. Physical cues, on the other hand, are micro- or nanolevel topographical modifications of the surface of the material, including surface roughness [3]. The most widely known example of response to topographical cues is the cells’ ability to recognize topographical differences, such as presence of ridges and grooves, and their tendency to be guided by these features [4]. The effect of topography is not restricted to alignment, it also affects adhesion, proliferation, ∗ To whom correspondence should be addressed. Tel.: (90-312) 210-5180. Fax: (90-312) 210-1542. E-mail: vhasirci@metu.edu.tr 1242 V. Hasirci et al. overall morphology and gene expression pattern (i.e., phenotypic effects such as differentiation) [5]. Thus, knowledge of the impact of these features on cell behavior is essential in both basic cell biology research and in more direct application-related areas, like tissue-engineering and cell/device-interface-based technologies. In tissue engineering, adhesion of cells to a cell carrier is the first step towards the development of an artificial tissue. Natural polymers, such as collagen, fibrinogen and elastin, are superior in this sense over synthetic polymers since they inherently possess specific amino-acid sequences that can be recognized by the cells. Conventional scaffold manufacturing techniques, such as solvent casting, freeze drying and salt leaching, produce scaffolds in which individual polymers are randomly oriented [6]. Thus, even though the carriers may be able to induce adhesion and proliferation of cells, cell orientation on these carriers would remain random. This would present an obstacle in certain tissues where tissue function is directly related with cell orientation. Two examples of such tissues can be corneal stroma and peripheral nervous system. In the natural corneal stroma, cells inhabit an environment which is composed of aligned fibrils of collagen. This specific orientation is shown to be essential for the transparency of the cornea [7] and an artificial cornea should mimic this. In the latter case, it is of utmost importance for a nerve guidance conduit to direct newly outgrowing neurites in a given direction for closing a gap within the nervous system. It is apparent that for these cases (and also some others like artificial tendon structures) a functional artificial tissue is not possible without controlled carrier topography that would guide the orientation of cells. For nerve guidance conduits, this can be achieved in 2D or 3D, by chemical or physical patterning of the surface by adhesion proteins, and by grooves and ridges, respectively. Their presence can guide the movement of the neurites in the desired direction [8]. For an artificial corneal stroma, surfaces modified by microor nanopatterns that imitate the natural state can be created. Patterning techniques and nanopatterns There are different methods to modify surfaces and provide patterns to achieve organization of cells. These patterns could be 2D or 3D, although the resulting cell organization is generally in 2D. The dimension of the patterns could be at the micro- or nanometer level. Achieving nanopatterns is a more difficult procedure than micropatterns because low micron level is what can be achieved with most of the current techniques employed. Studies on the influence of micropatterns on cell guidance (Fig. 1) are showing an increase and the data accumulated indicate that cells could be responsive to nano-level chemical and physical cues. Some of the methods developed to create nanopatterns are atomic force microscopy (AFM)based techniques (dip-pen nanolithography), hot embossing lithography and soft lithography. Nanobiomaterials: review of the existing science and technology, and new approaches 1243 (a) (b) (c) Figure 1. (a) Collagen film with micropatterns prepared by lithography. Shown is the nanolevel collagen organization corresponding to the ridges of the ridge–valley-type pattern. (b) Retinal pigment epithelial cells aligned on polyester blend micropatterns prepared by lithography with added fibronectin cues on the surface (day 1, Acridine orange staining). Shown is alignment and restriction of cell distribution by the patterns. (c) Retinal pigment epithelial cells aligned on polyester blend micropatterns prepared by lithography with fibronectin cues on the surface (Day 10, Acridine orange staining). Aligned cells within the patterns and disorganized cells at the unpatterned edges (darker strip at the top) can be seen. 1244 V. Hasirci et al. Dip-pen nanolithography In this method, mica wafers are coated with gold using thermal evaporation. Lithography is performed by using an AFM in its tapping mode. Contact mode could not be used, since the molecules transferred are generally proteins and they are sensitive to shear and frictional forces created by the application of this mode. AFM tips are first dipped into a protein (or polymer solution of choice), then the tip is tapped on the surface in a predetermined manner creating patterns on the surface. By this method, protein molecules have been patterned in nanometer thick lines, with line widths as small as 30 to 50 nm [9]. This technique is essentially ideal for patterning of biological molecules because the conditions of the method are milder in comparison to techniques like ion-beam-based lithography (can cause denaturation). Hot embossing lithography This technique enables us to achieve the production of topographical nanostructures at a scale of 10 nm laterally, with high throughput and at low cost. A thermoplastic film substrate is embossed by application of pressure and heat. By demolding and reactive ion etching, windows were opened to substrate after which silane deposition was done [10]. Silanes were used, since they can covalently bind onto surfaces. After lift-off by a solvent like acetone, a structured silane layer was obtained. Selfassembled monolayers, as well as functional biomolecules, can be used as coatings for use in biomedical applications. For immobilization of such molecules, the head group of silane should be modified in a way that it allows covalent bonding. These coatings could be done in a variety of ways including vapor deposition, spraying, casting and immersion. Soft lithography In soft lithography, the pattern is first created on a silicon wafer by use of a photoresist and standard photolithography. This pattern is then used as a template for making an elastomeric inverse replica of it from PDMS. This patterned stamp is immersed into the solution of the molecule that is used to create the desired final pattern and then pressed against the surface to transfer the molecules. Since the solvents used in this method are mild ones, bioactive molecules can also be patterned [11]. PDMS used in soft lithography is good for micro-scale patterns but could not be used efficiently in high-resolution cases like nanopatterning because of its low modulus. Therefore, researchers used composite materials or hard PDMS (h-PDMS) which has a higher tensile modulus and obtained stamps or molds to use in further steps of soft lithography. These stamps had various nanometer-scale patterns including sub-100 nm patterns [12]. This method was further developed to obtain patterns with feature sizes as small as 30 nm [13]. There are a number of successful applications of chemical and physical patterning at the micro-scale. Investigations have shown a positive influence of patterned Nanobiomaterials: review of the existing science and technology, and new approaches 1245 surfaces on the proliferation of cells. Cells were shown to align along the grooves and assume a more elongated appearance, in contrast to the cells seeded on smooth surfaces, which stay in a round form. Spreading is important for cell division; thus, appropriately patterned surfaces facilitate cell division by providing a surface on which spreading of cells is encouraged. More subtle changes related to the presence of patterns like the increased cytoskeletal and focal adhesion complex organization along the direction of the patterns have also been observed [14]. Moreover, it is speculated that this orientation may induce the oriented secretion of newly formed extracellular matrix molecules [15] and, thus, in a way improve the remodeling process. In addition, surface patterning has also been shown to increase some synthesis pathways for certain cell lines, such as increase in mineral production by osteoblasts on patterned surfaces [16] or increased alkaline phosphatase activity of osteoblasts [17]. Recent investigations have questioned whether effect of surface anisotropies at nano-scale has similar or superior effects [18]. Since nano-scale features are much smaller than the cell dimensions, their effect could mainly be on the formation of focal adhesions or single surface binding membrane proteins. Thus, such surface patterns might be more effective in precise control of cell directionality and migration than unpatterned surfaces. Responsiveness to nano-scale topography has been shown with several cell types, such as meningeal cells [19], corneal epithelial cells [20] and fibroblasts [21]. Moreover, these have demonstrated that different cell types respond to the same topography in markedly different ways. Overall reaction to the patterns remains the same; however, each cell type has its own limits of recognition and their own preferential topographical dimensions. Hence, we are in the dark in this field about the optimization of the surface topography with respect to cell types. Such research would definitely contribute to the design of novel tissue engineering scaffolds that take into account the optimal topography for the cell type used. Surface patterning designs are not restricted to ridges and grooves. There has been a growing amount of data concerning the effect of other types of features, such as curvature and presence of discreteness. Several groups have shown that cells respond to presence of nanopillars or pits and these topographical features adversely affect cell adhesion [22, 23]. This might be very promising, since in certain implants, such as stents, the ability to deter cells from adhering to the surface is an advantage. This is one of the areas where nanopatterns are more beneficial than micropatterns. An important parameter in this kind of topography is the spacing of the patterns and the height or depth of the patterns. It was shown that shorter pillars (or islands) improved the fibroblastic cell adhesion and spreading [24]. Again optimal dimension for effectiveness becomes important since island size was shown to significantly alter cell behavior. Other data involving cell response to topography stems from aligned nano- and micro-fiber experimentations [22]. Lastly, there have been attempts for nano-scale patterning of proteins on a surface in correct orientation for facilitation of cell adhesion [9]. 1246 V. Hasirci et al. Contact guidance is dependent on several factors. For example, the ability of filopodia and lamelopodia to recognize surface topography and to direct the cell spreading and movement accordingly is one of the foremost reasons that cells orient themselves with respect to the topography [25, 26]. Another important parameter is the elevation, where cells generally align along the grooves of micropatterned surfaces, because they cannot climb very steep surfaces. For negative effects reported with nano-scale obstacles it was hypothesized that presence of these prevents the formation of focal contacts, thus inhibiting the spreading of the cells. As mentioned above, height and spacing of nano-level surface modifications is a crucial factor in determining the behavior of cells. Other factors that must be considered are the effect of topography on protein absorption and possible chemical changes on the surface during pattern formation. Thus, not only the geometry and dimensions of the patterns are important, the pattern formation method must also be taken into account. More importantly, with the development of precise 3D tissue engineering scaffold production techniques, this information should be translated into 3D forms [27]. Earlier research in our laboratory have dealt with contact guidance on micropatterned natural and synthetic polymeric systems, such as collagen, poly(L-lactide) (PLLA) and poly(hydroxybutyrate-co-hydroxyvalerate) (PHBV) films [17, 28]. Our current efforts are focused on the possibility of aligning collagen fibrils along a microscale template, thus achieving nano-scale orientation on the scaffold. By using human corneal keratocytes as model cell system, the effect of micropatterning on proliferation, phenotype perseverance and morphology of keratocytes has been investigated. Studies were also conducted on contribution of patterning on the mechanical properties of the cell carrier in the presence of cells. Among further aims are the investigation of the effect of patterning on ECM secretion by keratocytes and observation of whether the presence of the patterns prompt keratocytes to secrete oriented ECM that would lead to transparency of the carrier after the removal of the original scaffold by degradation. Study of positive or negative effects of nano-scale patterning, either by fibrillation of collagen or nanolithography, on these parameters are being planned. Along with the ongoing experiments with nano-scale aligned fibrils, we also aim at contributing to the accumulating information about the responses of cells to different nano-scale topographies. FIBERS AS NANOBIOMATERIALS An important class of nanomaterials on which intensive research has been carried out in the last decade is polymeric fibers, which may be either at the micro- or nano-scale (Fig. 2A and 2B, respectively). Among these, nano-scale-sized fibers have attracted more attention in the last decade. The word nanofiber generally refers to a fiber having a diameter less than 100 nm, but by definition, fibers with diameters less than 1000 nm produced via some ultrafine manufacturing techniques such as electrospinning are also classified as nanofibers. Nano-scale Nanobiomaterials: review of the existing science and technology, and new approaches 1247 (A) (B) Figure 2. (A) Fibers produced by wet spinning of PHBV. (B) Fibers produced by electrospinning of PHBV. properties of fibers, especially resulting from the high surface area to volume ratio, make them preferable in many industrial applications [29 –31]. Research is mainly concentrated on the fabrication methods such as electrospinning, phase separation, drawing, template synthesis, melt-blowing and self-assembly [29, 30]. Among these, electrospinning is the most popular and preferred technique since this process is economical, simple, yields continuous fibers (while the others are a few micrometers in length) and is versatile enough to be applied to a variety of materials. While electrospinning results in fibers with a diameter in the range from 3 nm to several micrometers, other methods such as self-assembly, template synthesis and phase separation produce fibers with diameters ranging from 500 nm up to a few micrometers and their fibers are only a few micrometers long [30]. Thus, electrospinning is the most extensively used fabrication method. Due to their intrinsic features polymeric nanofibers are attractive for many practical applications. Biomedical and biotechnological applications such as tissue engineering, nanocomposites for dental application, controlled drug delivery, medical implants, wound dressings, biosensors and filtration are among the most intensively studied areas. Advantages of using such nano-scaled fibers in many industrial applications are mentioned in a great number of studies. Below, a review on biomedical applications of polymeric nanofibers is presented. Tissue engineering The design of an ideal tissue-engineering scaffold is one of the most important challenges. This scaffold should not only mimic the structure and biological functions of extracellular matrix (ECM), but should also provide a good environment for the cells so that they can easily attach, proliferate and differentiate. Human cells are known to attach, grow and organize well on fibers with diameters smaller than that 1248 V. Hasirci et al. of cells [18]. Polymeric nanofibers are of great interest because they can serve as tissue-engineering scaffolds and their dimension and properties arising from their dimension make them interesting carriers. Research is, therefore, concentrated on nanofiber applications in such areas as cartilage, nerve, bone, skin, skeletal muscle and blood vessel tissue engineering [24, 32 –36]. Among the biocompatible and biodegradable polymeric materials that have been used in the fabrication of scaffolds made from nanofibers are PLLA, poly(lactic acid-co-glycolic acid) (PLGA), poly(ε-caprolactone) (PCL), PHBV and a variety of their blends. Bone tissue engineering is one of the most attractive areas. Studies involving the use of PHBV and the effect of fiber thickness on cell behavior are being investigated in our group. Bhatarai and his co-workers used electrospun chitosan nanofibers in studying the behavior of chondrocytes and osteoblasts [37]. It was observed that adhesion of cells was promoted and cell morphology was good, showing the potential of these nanofibers as good carriers in bone tissue engineering. Electrospun nano/microstructured PLGA-based scaffolds have been shown to provide both guidance and flexibility to cardiac myocytes [24]. In another study, a self-assembled peptide nanofiber scaffold was used to create a basement membrane. The tailor-made peptide scaffold showed the high signaling capacity to enhance the formation of confluent cell monolayers of human aortic endothelial cells [38]. What remains a problem in electrospinning is the alignment of the nanofibers. If this is achieved successfully, the arrangement of cells in 2- and 3D architectures could be controlled to achieve better cell proliferation, differentiation and functional longevity. This was shown by Xu et al. [32], who used poly(L-lactic acid-co-εcaprolactone) (75:25) block co-polymer as a scaffold for blood vessel engineering. They showed that smooth muscle cells could attach and migrate in the direction of aligned nanofibers and project spindles indicating the suitability of the scaffold for artificial blood vessel. Wound dressings There is an increasing demand for advanced wound care products and, thus, for new wound-dressing materials. Absorption of exudates, ability to provide and maintain a moist environment, to adhere specifically to healthy tissues, ease of removal without pain and low cost are some of the requirements of an ideal wounddressing candidate. Normally these products are made from hydrogel sheets or hydrophilic, microfibrous fabrics. The use of polymeric nanofibers in this context is new. Electrospun biodegradable nanofibers are directly sprayed onto the injured part of skin. The fibrous mat dressing created on the skin encourages the growth of normal skin. In this way, scar tissue formation, which occurs in normal treatments, is eliminated [29]. Virginia Commonwealth University researchers had success in creating a nanofiber mat that would act as a ‘natural bandage’ [39]. When electrospun collagen nanofibers were placed on the wound, blood loss was eliminated and the normal wound-healing process was successfully achieved. Nanobiomaterials: review of the existing science and technology, and new approaches 1249 Recently Rho et al. [40] investigated the potential of a biomimetic nanofibrous extracellular matrix in tissue engineering and wound healing. Two rectangularshaped (1 cm × 1 cm) wounds were created at the back of rats and cross-linked collagen nanofibers (diameter range 100–1200 nm) were applied on the wounds. The wounds treated with cross-linked fibers exhibited an early-stage healing when compared to the cotton gauze controls. Implants There are some US patents that have been issued for fabrication techniques of polymeric nanofibers for their use in vascular and breast prosthesis applications [41, 42]. Other implant uses have been exploited over time. In a study, electrospun, nonwoven and bioabsorbable PLGA anti-adhesion membrane was impregnated with an antibiotic for an in vivo rat model and the results showed a reduced post-surgery adhesion, indicating that the nanofibrous membrane could both be used as a local drug delivery vehicle and a physical barrier [43]. In addition to this, Buchko et al. did research on biodegradable porous protein polymeric films [44]. Silicon substrates were coated with a thin nanofibrous polymer film, which was then used as a model for a neural prosthetic device. The aim of using this device was to investigate its capability of recording and stimulating the neural signals that occur in case of neural damages. This coating worked as an interphase between the prosthetic device and the neural system by reducing the stiffness mismatch between these two phases and, thus, preventing device failure after implantation. Controlled release systems Delivery of drugs at the appropriate time, duration and site are among the most important points that have to be considered while designing controlled release systems. Also, the biocompatibility, solubility and stability of the drug in the body are important. Some drugs may be poorly soluble in water, while others may exhibit poor stability, high toxicity and low bioavailability if administered as is. Therefore, there is a need for drug carriers that prevent direct exposure of the bioactive agent to in vivo medium and modify the availability of the drug as required. This carrier should be of appropriate dimension (<100 nm) in order to escape filtration by capillary beds, should have sufficiently high molecular weight (>70 kg/mol) to avoid renal extraction, should be biocompatible and should provide appropriate release kinetics. Polymeric nanofibers are gradually taking their place among the drug-delivery vehicles. Their large surface area to volume ratio and their nanometer scale aid mass transfer and could provide sufficient drug release. The drug of choice is generally incorporated into the polymer by mixing before electrospinning and depending on the properties of the drug the resulting nanofibrous structure may have different forms. The drug spun together with the polymeric solution may be localized on the 1250 V. Hasirci et al. surface of the nanofibers or a homogeneous blend of both the carrier and drug can occur [30]. The advantage of using this nanofibrous carrier type is that site-specific delivery from the scaffold into the body can be obtained by implantation via surgical interference and the method is applicable to both hydrophilic and hydrophobic drugs. Prevention of post-operative infections and adhesions are two major concerns in medicine. A hydrophilic antibiotic, cefoxitin sodium, has been incorporated into electrospun PLGA-based polymeric micelles. Its ability to prevent infections that appear after surgery and release behavior was investigated [45]. First of all, it was observed that the amount of drug incorporated into the fibrous scaffolds influenced the density (a decrease was observed) and the morphology (bead size was reduced as concentration of drug was increased). This was concluded to occur as a result of salt increase in the structure with increase in the drug amount. Secondly, it has been observed that the release profile was more controlled in the initial stages of release and was maintained for longer durations (1 week more) in the case of PLGA blend (PLGA/PLA/PEG-b-PLA, 80:5:15) when compared to drug incorporated-PLGA nanofibrous scaffolds. This was due to the presence of the hydrophilic block copolymer (PEG-b-PLA), which may have encapsulated a certain portion of the drug inside it rather than leaving it on the surface of the scaffold. Another conclusion was that electrospinning process did not affect the structure of incorporated cefoxitin sodium drug, since the drug maintained its inhibitory function on bacteria even in that form with the same efficiency as the pure drug. All these proved the potential of using nanofibrous PLGA blends in preventing post-surgical infections and adhesions. Another study, showed the efficiency of using electrospun fibers in wound healing and topical delivery [46]. In this case, the release of a poor water-soluble drug was studied. An amorphous nanodispersion of both ketanserin and itraconazole with the polymer polyurethane was obtained. They concluded that water insoluble nonbiodegradable polymers could be a good choice for local delivery of insoluble or poorly water soluble drugs or in wound healing. This shows the potential of electrospun nonbiodegradable polymeric nanofibers in solving the problem of delivering poorly water-soluble drugs in a controlled manner. Reinforcement of composites Micro- and nano-sized fibers can be used to reinforce the composite structures. Since the properties of materials change as their size is reduced to nano-scale, nanofiber-reinforced composites are expected to have superior properties in comparison to the traditionally used composites. Certain applications are currently under development. These studies mostly concentrate on composites reinforced with either carbon nanotubes or nanofibers obtained by processes other than electrospinning [29]. Vapor-grown carbon nanofibers (VGCF) have high mechanical and physical properties and are being increasingly used in improving mechanical, thermal and electrical properties of polymer-based materials. In a study, carbon nanofibers Nanobiomaterials: review of the existing science and technology, and new approaches 1251 (CNF) were used to reinforce poly(ether ether ketone) (PEEK) structures to obtain nanocomposites with superior properties [47]. Injection molding and extrusion are the main polymer processing techniques used to obtain high-temperature semicrystalline polymer nanocomposites. In this case, different concentrations of CNFs, namely 0, 5, 10 and 15%, were incorporated into the polymer by means of an extruder. Nanocomposite mechanical properties such as tensile strength, Young’s modulus, yield stress and strain were found to increase with increase in the CNF fraction. Furthermore, CNFs were observed to be homogeneously introduced and aligned in the thermoplastic composite, indicating a good interaction between the matrix and CNF fillers. Another important application of polymeric nanofibers as reinforcement materials is the production of orthodontic composites, where they could be used as fillers [48, 49]. Here also the mechanical properties like elastic modulus, flexural and tensile strengths and stiffness has been successfully improved. Traditional dental amalgams are made up of a resin matrix and a filler; in the recent studies they are being replaced with polymeric restorative composites. A variety of fillers are used in varying concentrations (depending on their characteristics) to formulate these composites. The properties of these polymer-based composites have been shown to be similar to those of traditionally used ones, but the inorganic filler particles contributed to the failure of the composites and led to short service lives (generally 12–18 months). Thus, there is a need for another type of filler that will increase both the mechanical properties and the service life. Fong [48] has used electrospun Nylon 6 nanofibers in place of particulate fillers in the formulation of the dental methacrylate, 2,2′ -bis-[4-(methacrylopropoxy)-phenyl]propane/tri(ethylene glycol) dimethacrylate (BIS-GMA/TEGDMA) resin. The results showed a significant increase in flexural strength (36%), elastic modulus (26%) and work of fracture (42%) of the nanofiber-reinforced composite resins even with the addition of very small amounts of Nylon 6 nanofibers (5%). A rougher surface structure and less fracture steps were observed on the reinforced composite as compared to the neat resin sample, which show that nanofiber could deflect the crack. An increase in the resistance of material to fractures was observed; thus, the fact that even small amounts of Nylon 6 nanofiber filler could enhance the properties of dental composite shows clearly the importance of using such fillers. It should be mentioned that, there is still work to be done in this aspect. The fact that until now the electrospinning process did not produce sufficient amounts of aligned nanofibers and continuity is a limitation against their widespread use as reinforcements in composites [30]. Despite all of these, the studies carried out until now have indicated that polymeric nanofibers could be the ideal future components of orthodontic composite applications. Filtration Use of polymeric nanofibers in filtration systems such as gas, liquid and molecular filtrations, is another important biomedical area to be considered. Electrospun 1252 V. Hasirci et al. nanofibrous membranes are among the most preferred filtration materials due to the fact that filtration efficiency increases with decrease in the diameter of fibers. Other reasons for using polymeric nanofibers in filtration are their lightweight, flexibility and nanometer size of the pores (which in the traditional ones vary from few micrometers to several micrometers, depending on the application). Ahn and co-workers [50] proved the high efficiency of electrospun Nylon 6 nanofibers in filtering test particles with a diameter of 300 nm at a face velocity of 5 cm/s. Despite high-pressure drops across the nanofilters, they were accepted to possess a high capacity for filtration. Enhancement of filter life in pulse-clean cartridge applications for dust collection and increase in the efficiency of air filters for personnel cabins of mining vehicles are some other possible applications of polymeric nanofibers [51]. Polymeric nanofibers are ideal materials for molecular separation because of their high surface area to weight ratio. As a result of this property, polymeric nanofibers can be used as the stationary phase on which a specific molecule can be immobilized by a variety of surface modification methods. The immobilized molecule can then interact with the specific solute and achieve separation. The immobilized molecule could be an antibody for a specific protein and separation of that protein from a mixture of proteins can be achieved successfully [29, 30]. Ma et al. have investigated the use of electrospun cellulose acetate nanofiber membranes as affinity membranes [52]. Higher water permeability of nanofibrous membrane when compared to traditional microporous membranes was observed. Moreover, it has been mentioned that membrane could be regenerated by rinsing with elution buffer, which makes the membrane reusable. In this kind of applications, it is important to choose a material that does not possess high specific protein adsorption. As a consequence, hydrophilic polymers are preferred more than the hydrophobic ones. Biosensors There is widespread use of biosensors among which food, environmental and clinical applications can be counted. A biosensor is composed of a transducer and a biofunctional membrane and its function is to convert a biological signal into an electrical output. The sensing membrane should be carefully chosen since factors like aging, reproducibility, response time, selectivity and sensitivity are all dependent on it. Polymeric nanofibers are attracting attention in biosensor applications again due to their large surface area to weight ratio. This characteristic is crucial because in many cases detection of substances with very low concentrations is required, and this can be done only with materials possessing high sensitivity [30]. Since enzymes are abundantly found in the nature and are highly specific to their substrates, they are very suitable for use in the area of biosensors. A good example of a biosensor is the urea biosensor proposed by Sawicka et al. [53]. Electrospinning has been used to obtain a non-woven mat of biocomposite nanofibers (consisting of a mixture of urease dissolved in buffer (30%) and the poly(vinylpyrrolidone) Nanobiomaterials: review of the existing science and technology, and new approaches 1253 polymer dissolved in ethanol (70%)). It was observed that the enzyme retained its activity in the polymer solution and what is more important is that the enzyme performed its catalytic functions, even in small concentrations and the response time was short. In another application, a composite of Au nanoparticles and conductive polyaniline nanofibers has been proposed to be a good base for designing a glucose biosensor. Glucose oxidase was immobilized on the surface of nanofibers and then was used in the detection of glucose concentration. The nanocomposite biosensor exhibited high reproducibility and stability and very good glucose detection performance [54]. NANOTECHNOLOGY AND DRUG DELIVERY The recent developments in the field of nanotechnology drastically improved the area of nanomedicine. There are several benefits of nanotechnology-based approaches to therapy, personalized medicine, intelligent drug design and targeted drug delivery. Personalized or individual-based medicine aims to develop drugs according to the patient’s genotype by making use of nanoarrays for molecular diagnostics, whereas conventional approach tries to match the existing drugs with the patients in the most suitable way [55]. Intelligent drugs are being developed to respond to stimuli and specifically react with the target and existing drugs are being modified so that their side effects, immunogenicity or toxicity might be decreased. Nanotechnology approaches can also be used to augment the effectiveness of the molecule as a therapeutic agent. For example, the poor aqueous solubility of drug candidates limits their bioavailability and the drug-discovery process. This solubility limitation can be addressed by reducing the drug particle size to nanometer scale [56]. Controlled drug-delivery systems are developed to maintain appropriate doses locally for prolonged periods. Targeted drug delivery aims to transport active agents to predetermined locations in the body, increasing the efficacy of the agent, while decreasing the systemic side effects compared to conventional administration routes. When drugs or other bioactive agents are delivered by a controlled release system, the observed increase in the activity of the agent is due to protection of the active agent against degradation within the carrier. Targeting, on the other hand, decreases the amount of bioactive agent needed. Such systems also help mask unpleasant taste of certain drugs [57]. Most of these aims have been achieved with conventional delivery systems [58, 59]; however, with such systems, intracellular delivery and delivery across some physiological barriers is not possible. The promise of utilization of nanotechnology in drug-delivery systems is the possibility of intracellular delivery of agents such as DNA [60], anti-sense RNA and anticancer drugs [61], as well as penetration through barriers such as the blood–brain barrier [62] and tight junctions, therefore permitting the accumulation of agent at previously ‘unreachable’ target sites. Nanoparticles of all sorts can penetrate through capillaries, their uptake by the cells and movement through the dense 1254 V. Hasirci et al. extracellular matrix is evidently easier than with microparticles. This has another advantage; since penetration capability of nanoparticles is high, local delivery through injection can be achieved with a minimal damage to the tissue, since introducing the nanoparticles in close proximity to the site is sufficient. Moreover, utilization of nano-scale carriers would increase the control over drug dosage which enables directing small amount of drugs to a desired site. Enhanced targeting decreases the total amount of drug used, which, in future, may reflect to the market as a total decrease in drug prices. Increased efficiency of nanoparticles due to their small size can most convincingly be shown by the accumulation at the tumor site. Since vasculature at the tumor site develops very fast, they tend to be leaky. It was shown that the pores of the capillaries have dimensions around 100–1000 nm, in contrast to the pore size of 10 nm observed in healthy tissue [63]. Thus, the ability to prepare delivery systems within this range would ensure the accumulation at tumor sites due to enhanced permeability and retention effect. Since carriers cannot escape the circulation at other sites, they certainly would not accumulate within the veins because of their size. In addition, particles less than 100 nm in diameter can move through the pulmonary system with more ease [66], while particles around 100 nm were better absorbed in the gastrointestinal tract [64]. Nano-scale designs are especially important for systems to deliver drugs with low therapeutic indices. Since control over drug release and targeting is more precise in such systems, toxicity can be averted. Nano-scale drug-delivery systems constructed of polymers and ceramics could also be made to deliver with a zero-order behavior, so that constant levels of drug within the body can be obtained throughout the course of application [63]. There are several forms that can serve as drug carriers at the nano-scale. The most obvious controlled release system forms are microspheres, microcapsules and liposomes, which are scaled-down to nanometer range by improved manufacturing methods. Micelles and dendrimers are also other important controlled release system forms currently receiving great deal of attention. The details of these approaches are provided below. Nanoparticles in drug delivery Solid, colloidal substances varying in size from 10 to 1000 nm are defined as nanoparticles [66]. An active agent is coupled with the nanoparticle through entrapment, adsorption, attachment, encapsulation or directly dissolving the agent within the nanoparticle structure. There are several methods to form nanoparticles: these are molecular self-assembly [67], nanomanipulation [68], photochemical patterning [69] and bioaggregation [70]. These lead to solid or hollow nanospheres, porous or solid nanoparticles, depending on the method of preparation. Nanospheres and nanocapsules are the two most widely employed structures and have different properties and release characteristics for the therapeutic agent carried. Nanospheres are solid particles in which the active agent is physically and uniformly dispersed, Nanobiomaterials: review of the existing science and technology, and new approaches 1255 whereas nanocapsules are hollow particles where the content is encapsulated by a membrane [71]. Effectiveness of delivery can be increased by targeting. There are two possible ways of achieving targeted delivery, i.e., active and passive targeting. Active targeting is done by attaching the active agent or the carrier system to a tissueor cell-specific ligand [72], whereas passive targeting involves coupling the active agent to a macromolecule such as a high-molecular-weight polymer that passively reaches the target organ [73]. Nano-scale particulate production may also open the road for novel ceramic-based delivery systems. Ceramics such as silica, alumina and titania are known to be biocompatible [74], but their hardness relative to the natural tissue restricts their usage to only hard-tissue components. However, at nano-scale, their mechanical abrasion capability is less pronounced. They have several advantages over their polymeric counterparts, such as higher stability under different pH and temperature conditions and better protection of labile agents against denaturation [74]. Surface modification with functional groups is possible and this enables conjugation with different antibodies or ligands leading to effective targeting in the body [75]. Current research in our laboratory includes preparation of nanocapsules of PHBV for the delivery of growth factors targeted to bone. The nanocapsules are prepared by the water-in-oil-in-water technique. Among the investigated production parameters are the concentration of PHBV, nature of the organic phase and the surfactant, all of which significantly affected the morphology of nanocapsules as shown in Fig. 3. Micelles in drug delivery Micelles are structures formed by co-polymers in which the hydrophilic part of the structure faces the outer aqueous environment, while the hydrophobic structure constitutes the inner part; therefore, they are excellent candidates for the delivery of water-insoluble drugs. Polymeric micelles are thermodynamically more stable than surfactant micelles and this contributes to the increased stability of the drug to be administered [76]. Recently, Wang et al. [77] have shown that the poorly water-soluble anti-cancer drug Paclitaxel can be successfully carried within mixed polymeric micelles consisting of poly(ethylene glycol)-distearoyl phosphoethanolamine conjugates (PEG-PE), solid triglycerides (ST) and cationic Lipofectin lipids (LL). The micelles prepared had an average size of 100 nm and were shown to be stable that no drug release was observed during 4 months of storage. In vitro anti-cancer effects of PEG-PE/ST/LL/paclitaxel and control micelles were tested on various cancer cell lines, and paclitaxel in PEG-PE/ST/LL micelles demonstrated the maximum anti-cancer activity. The size of polymeric micelles (approx. 100 nm in diameter) provides them with a way to distribute their contents more efficiently and also gives the increased ability of intracellular delivery and penetration through endothelial lining which is an essential feature for reaching deep tumors. Also their renal clearance is slower 1256 V. Hasirci et al. Figure 3. PHBV nanocapsules formed by water-in-oil-in-water technique. (a) Continuous medium: 1%(v/v) Tween-20 in PBS (pH 7.4), PHBV: 120 mg. (b) Continuous medium: Tris-HCl buffer (pH 7.4), PHBV: 30 mg. because of their size and hydrophilicity, which provides them the ability of long circulation in vivo [78, 79]. Polymeric micelles are also useful as targeted drug-delivery agents because they can show stimuli-responsive behavior. It was shown that polymeric micelles based on PLLA-poly(2-ethyl-2-oxazoline)-PLLA(PLLA-PEOz-PLLA) ABA tri-block copolymers achieve intracellular delivery of the anti-cancer drug doxorubicin and change micellar structure with the change in intracellular pH [80]. Therefore, drug release from micelles was inhibited at pH 7.4, whereas accelerated release was observed at acidic conditions, leading to selective destruction of cancer cells. Nanobiomaterials: review of the existing science and technology, and new approaches 1257 Dendrimers in drug delivery Dendrimers are macromolecular compounds which are comprised of an inner core surrounded by branches extending outwards [81]. Dendrimers can be synthesized starting from the core (divergent synthesis) or starting from the outermost branches (convergent synthesis). Rather than being randomly branched macromolecules dendrimers are highly controlled structures. This complexity endows them with several structural advantages such as naturally being in nanometer size, ease of production and functionalization [82]. In the study of Patri et al. [83], to reduce toxicity of amine-terminated dendrimers and to increase aqueous solubility, surface hydroxyl groups were modified to have a neutral terminal functionality for use with surface-conjugated folic acid as the targeting agent. It is also possible to functionalize a single dendrimer with several different groups which might increase its effectivity. Like in the micelles, dendrimer cores and branches can be prepared separately, i.e., a hydrophobic core can be made to interact with hydrophilic branches. The presence of internal cavities within the dendrimers makes them suitable candidates for drug carrying. Drugs can be loaded within the central core portion or at the branch sites either by encapsulation or complexation. The highly branched structure of the dendrimers is a concern, since it generally induces immune response [84]; however, this property can be exploited in design of vaccines [85]. INTELLIGENT SYSTEMS BASED ON SMART NANOBIOMATERIALS One of the most promising applications of nanotechnology involves design and development of intelligent delivery systems which are capable of showing responsive behavior upon a certain environmental signal such as temperature, pH, ionic strength, electric and magnetic field. Nano-scale responsive delivery systems are very promising, since they offer a number of advantages such as specific targeting, stimuli-dependent release behavior and enhanced ability of escaping from phagocytotic uptake and, thus, prolonged circulation times owing to their nano-scale sizes [80]. Many studies have been conducted on delivery of protein-peptide drugs, genes and anti-sense oligonucleotides via such intelligent nanosystems [86 –89]. The stimuli sensitive characteristics of certain natural or synthetic polymers which can undergo fast and reversible changes upon small environmental changes are used for designing a responsive system. These ‘smart’ polymers can be used in various physical forms including, micelles, crosslinked (permanently) hydrogels, reversible hydrogels, modified and conjugated solutions [90 –92]. The main mechanism underlying such a system is the conformational change triggered by an environmental change which causes a loss of integrity of the nanobody, resulting in the release of the bioactive agent entrapped within it or bound to it. The most extensively investigated nano-scale responsive delivery systems include temperature, pH and magnetic field responsive systems. 1258 V. Hasirci et al. Thermoresponsive systems Temperature-sensitive nanosystems for controlled delivery are being widely studied. Certain polymers have the ability of undergoing phase transition in a temperaturedependent manner; above or below a specific temperature they are water soluble and lose their solubility at temperatures below or above this temperature. These critical temperatures could either be lower critical solution temperature (LCST) or upper critical solution temperature (UCST). For LCST, solubility decreases with increase in temperature and for UCST solubility increases with increase in temperature. The most commonly studied temperature-sensitive polymers in nanobiomaterialbased delivery applications display LCST. These polymers possess both hydrophilic and hydrophobic groups in structure, below their critical temperature hydrophilic interactions dominate so they are water-soluble; however, above LCST hydrophobic interactions begin to dominate and they become water-insoluble. When in hydrogel form these polymers swell rather than dissolve in water below their LCST. This property of smart hydrogels provides a switchable swelling–deswelling behavior which is used to initiate release from smart nanohydrogels. Normally, release rate of bioactive agents from nanohydrogels is high at values below LCST due to higher swelling and low at values above LCST due to deswelling [90]. Preparation of hybrid nanogels has enabled scientists to design nanohydrogel based delivery systems that display a positive thermo-responsive release profile, at temperatures above the LCST drug-release rates increase, whereas they decrease at values below LCST. An example of such a hybrid nanogel is based on interpenetrating networks of thermo-sensitive poly(N-isopropylacrylamide) (PNIPAAm) gels and tailored nanoporous silica. This design enables the diffusion of the drug through the porous channels of silica at temperatures higher than LCST as the shrinkage of PNIPAAm opens the pores and squeezes the drug into the channels [93]. The swelling–deswelling kinetics, the LCST and UCST of responsive nanosystems can be altered by controlling the relative amounts of hydrophobic groups to hydrophilic groups. In order to gain control over these properties many studies have generated novel temperature-sensitive nanobiomaterials constructed of co-polymers in form of blocks, grafts or branches assembled in micellar structures [94 –97]. Apart from their relative ratio to the hydrophilic groups, the nature of the hydrophobic moieties of these co-polymers dramatically affect the thermo-responsive behavior therefore the alteration of hydrophobic moieties for synthesizing co-polymer structures offers great control over the release kinetics. A study has investigated the effect of micellar hydrophobic inner core chemistry on the temperature-responsive release behavior of a model hydrophobic drug from block co-polymers ordered as core–shell micellar structures. The outer shell polymer assigned for thermoresponsive behavior was PNIPAAm, whereas the substituted hydrophobic core moieties were poly(butyl methacrylate) (PBMA) and polystyrene (PS). The temperature responsiveness of PNIPAAm–PBMA and PNIPAAm–PS core–shell micellar structures differed from each other dramatically. The PNIPAAm–PBMA released their content upon increase in temperature above the LCST value of PNIPAAm, whereas Nanobiomaterials: review of the existing science and technology, and new approaches 1259 PNIPAAm–PS micelles retained their integrity and did not release the loaded drug above the LCST of PNIPAAm [98]. Other parameters, such as the cross-linker amount, polymer concentration, loading doses and the type of drug, have been investigated for achieving control over the thermo-responsive behavior of such nanosystems [99]. The effect of the nature of the loaded drugs on the thermoresponsive behavior was investigated on poly(N-vinylcaprolactam) (PVCL) nanoparticles loaded with three different model drugs. The results have shown that the temperature dependent release behavior is affected by the loaded drug types [100]. The most widely studied temperature-sensitive polymer is PNIPAAm with a LCST of 32◦ C [90]. Many studies have generated co-polymers of NIPAAm for designing temperature-sensitive nanosystems [95, 98 –101]. For instance, block co-polymers of poly(N-isopropylacrylamide-b-methyl methacrylate) were used for preparing self-assembled micelles to act as a vehicle for delivery of antiinflammation drug prednisone acetate as the model drug. The study has shown that the in vitro release behavior of prednisone acetate was dramatically and reversibly influenced by the temperature changes owing to temperature-sensitive micellar shell structure [96]. One very important application of thermo-responsive delivery systems based on nanobiomaterials is targeted delivery of therapeutics to tumor sites via local heating. The targeted delivery is achieved due to formation of aggregates at locally heated sites owing to the insoluble aggregate formation of polymeric systems above their LCST. Such an approach is important for reducing the systemic effects of chemotherapeutic agents which are known to affect healthy cells, as well as cancerous cells. By this approach, the therapeutic agent is used efficiently with the tumor site being exposed to desired levels of drug while the sideeffects of a systemic delivery are reduced to minimal levels. Hyperthermia by itself is known to have cytotoxic effects on cancerous cells; such an approach combines the therapeutic effect of hyperthermia with chemotherapeutic agents. Many studies have been conducted to design novel systems for hyperthermia directed targeted delivery to tumor sites. The LCST of PNIPAAm and an artificial elastin-like polypeptide (ELP) were altered by the tailoring of the two polymers via adjusting their relative hydrophobicity. Their LCST was designed to be higher than the physiological temperature and close to the temperatures used in clinical hyperthermia to assure temperature-dependent phase transition. In vivo evaluation was carried out with human tumor implanted nude mice. The results have shown that for ELP, an approximately 2-fold increase in tumor localization was achieved when compared to a control group of the same polypeptide tailored to be insensitive to temperature changes in the range used for hyperthermia. Similar results, but with lower degree of accumulation, were obtained for PNIPAAm [102]. 1260 V. Hasirci et al. pH-responsive systems Certain natural or synthetic polymers, called polyelectrolytes, possess ionizable pendant acidic or basic groups in their structure. Depending on the pH of the medium, these groups are protonated or deprotonated, leading to swelling/deswelling of the system and serve as the driving force for the release process [103]. In the design of a pH-responsive delivery system the most important step is the selection of an appropriate polyelectrolyte with a pKa value close to the pH of the targeted release medium. However, if a specific polymer needs to be used for reasons other than its pKa , the pH responsiveness of the whole system could be changed by incorporation of other ionic and/or hydrophobic moieties into the polymer structure [104 –106]. The oral delivery route could be preferable in the treatment of many diseases, especially for chronic diseases such as diabetes mellitus, for which current therapy methods involve painful, routine injections. The harsh acidic environment of the stomach is very threatening for peptide and protein drugs. To avoid degradation of these drugs in the stomach and achieve their targeting to the intestinal tract for subsequent absorption, pH-sensitive nano-delivery systems are being devised. The basic idea is the use of weak polyacids whose functional groups are uncharged at acidic concentrations, resulting in a collapsed state of the polymer preventing release and, thus, protection of proteins while passing through the stomach. After reaching the neutral or slightly alkaline intestines, the functional groups of these polyacids are deprotonated, negatively charged, leading to swelling and in targeted release at intestine [90, 107]. Nanospheres of cross-linked networks of methacrylic acid or acrylic acid grafted with PEG were designed for oral delivery of insulin. In vitro studies revealed that insulin was released at neutral pH, while it was retained in the nanospheres at acidic pH (pH 3.0). In vivo evaluation was also carried out and it was reported that the serum glucose levels were significantly lowered in diabetic rats that received the insulin-loaded nanospheres when compared to the control group [108]. Another use of pH-responsive biomaterials is non-viral genetic material delivery, which holds great promise for certain applications involving harvesting of therapeutic proteins via recombinant DNA technology and gene therapy for monogenic disorders or other states of disease such as cancer. The major obstacle encountered in non-viral genetic material delivery applications is the endosomal entrapment of the delivery vehicle. The enzymatic degradation of the vehicle and its contents before they reach the nucleus is one of the major reasons of low transfection efficiencies observed in non-viral delivery systems. In viral delivery systems displaying much higher transfection efficiencies the viral fusogenic peptides with pH-dependent membrane disruptive abilities overcome this obstacle. Certain polymers can mimic these fusogenic peptides. Weak polyacids with pKa close to the endosomal pH range (between 6.5 and 5.5 for early to late endosomes, respectively) undergo pH-dependent conformational changes after endosomal entrapment. These conformational changes allow insertion of polymer chains into endosomal bilayers Nanobiomaterials: review of the existing science and technology, and new approaches 1261 through hydrophobic interactions and cause their disruption allowing the release of contents into the cytoplasm [90]. Certain weak polybases, however, display another mode of pH-dependent endosomal membrane disruptive behavior. These polybases are protonated in the acidic endosomal environment and buffer the endosomal medium; this is followed by proton influx which leads to rapid osmotic swelling, membrane rupture and endosomal escape of vehicle contents. pH-responsive polyacids and polybases are employed in a variety of genetic material delivery designs [109, 110]. It was reported that when used as DNA carrier vehicles cellular uptake of nanoparticles smaller than 150 nm was enhanced by receptor-mediated endocytosis [111]. Combining this property of nano-sized DNA vehicle formulations with pH responsiveness to enable endosomal escape can yield novel non-viral genetic material delivery formulations with high transfection efficiencies. It has been reported that solid tumors display a lower extracellular pH when compared to normal tissue due to enhanced aerobic and anaerobic glycolysis in cancer cells [112]. Such nano-scale pH-responsive systems can be employed in specific targeting of tumors for cancer therapy [113, 114]. Magnetic-field-sensitive systems Dispersion of magnetite particles in polymer networks has generated magneticfield-responsive delivery systems [89]. The main application of these systems concern targeted delivery to tumor site via a locally applied magnetic field. When used in combination with temperature-sensitive polymers, it is possible to achieve not only specific targeting but also a temperature initiated release. Also, certain magnetite particles have the ability to generate heat as they are subjected to magnetic field [115]. The resulting hyperthermia triggers a conformational change in the temperature responsive component that directs the release of the bioactive agents. NANOCOMPOSITES IN HARD TISSUE IMPLANTS The rate of orthopedic implantation surgeries is experiencing a rapid growth. For example, the number of hip replacements has shown a 33% increase from 1990 to 2000 [116]. However, the cases of early failures are not uncommon and 10 to 15 years of lifetime is average for hip replacement implants [117]. This period is obviously not enough especially for young people and even more problematic when we think of the increased life expectancy in modern countries. Therefore, we must find a way to increase this service period of hard tissue implants. Complexity of mechanical behavior of natural bone tissues is a very important, but not the only, obstacle in designing properly functional yet durable biomaterials since bones provide the ultimate support to body. It is almost impossible to devise a monolithic (single-phase) artificial material that has both enough fracture toughness and is elastic enough to meet mechanical properties of bone. This led researchers to search 1262 V. Hasirci et al. for a composite (multi-phase) material that can approximate these mechanical needs in a bone replacement operation. With the recent advances in the nanotechnology field that allow creating and handling nano-sized particles and under the light of nature’s way of dealing with the issue, most of the studies in this field are conducted to find a proper design and use of nanocomposites to create an implant with sufficient mechanical properties. The term nanocomposite may be defined as heterogeneous combination of more than one material in which at least one of the phases must be in the nano-scale. The purpose of bone grafting studies is to design a scaffold which is able to restore the defects in bone by regeneration of the living tissues. Therefore, the integration of graft into the healthy bone tissue is a vital prerequisite of successful grafting in terms of both mechanical fixation and biological sealing. Different means of integration can be achieved by designing the graft in a way that interact with bone tissue in vivo in certain ways called osteoconductive, osteoinductive and osteogenic grafts. The first property is directly related with mechanical characteristics of graft, while the second and third ones are related with the capability of grafts to direct the cellular responses. Osteoconductivity of a graft serves as a template for host bone to infiltrate and regenerate throughout the defect site. Hydroxyapatite (HAp) is the most commonly used component in designing such grafts because of its high chemical and physical similarity to the mineral part of bone. Combining HAp with materials helps the integration of final composite biomaterial into the natural bone because the connection between bone and biomaterial is actually the sharing of apatite layers between them. Other bioactive ceramics are Bioglass® , β-tricalcium phosphate (TCP), HAp/TCP biphasic ceramics and glass-ceramic A-W [118]. Rationale for nano-scale structures Bone is a highly ordered assembly of macro- to nano-level hierarchical units, including cellular, organic and mineral constituents. Since bone is a nanocomposite itself, it can be expected to be used in the same architecture in a biomaterial that is designated as a bone graft. There are different calcified tissues in human body and their way of organization is different. All of these, however, have two things in common: the protein matrix component, collagen, and the inorganic, ceramic-like content called HAp. These two components are organized in nano-scale and therefore, bone itself is a true nanocomposite. Gao et al. studied the reasons for the superior strengths of natural nanocomposites from different origins like bone, tooth and nacre, all of which are made of nanocomposites of hard mineral plates or needles within a soft protein matrix, compared to an equivalently-sized monolithic structure of the same mineral [119]. They searched for a reason for the fact that the repeating subunits in these natural structures are all nano-sized. They assumed the proteins are, in practice, equivalent to cracks in a monolithic mineral crystal and deduced from their studies that there exists a critical length scale below which the fracture strength of Nanobiomaterials: review of the existing science and technology, and new approaches 1263 a cracked crystal is the same with that of a perfect crystal. This length scale is roughly 30 nm and lengths of mineral constituents of most hard tissues are around this number or they may be around up to ten times of this number probably because of some other design optimizations. Therefore, this is an important justification for studying composites in a nano-scale. Tissue response to nanocomposites Bone tissue has five distinct cell types that are responsible for creating, maintaining and remodeling the bone matrix with its unique hierarchical assembly. Two of these are especially important in determining the fate of grafted biomaterial: osteoblasts (bone forming cells) and osteoclasts (bone resorbing cells). The coordinated action of these two is crucial to maintain the health of the bone that forms around the implant. Therefore, scientists have concentrated on making implant surfaces attractive to osteoblast (osteogenicity) and favor their differentiation on the implant surface (osteoinduction). Indeed, Webster et al. [120, 121] have shown significant increase in protein absorption and osteoblast adhesion on nano-sized ceramic materials compared to µm-sized ceramic materials. These observations suggest the preference of nano-sized material (or surface) geometry by osteoblasts. The success of any bone implant is ultimately dependent on the osteogenicity of the graft, since assimilation of the new material into the bone tissue is only possible with the action of osteoblasts. Therefore, it is helpful to provide osteoblasts with a means of attachment and make them feel comfortable on this attachment site in this manner. This type of attachment is provided to some cell types with the presence of certain proteins such as fibronectin and vitronectin. These proteins are normally found in biological fluids and mediate adhesion, growth, and differentiation processes in a cell-type-specific way. Thus, if we know these specific cell–protein pairs and if we can get the proteins adsorbed on our biomaterial surface, then the desired cell types will automatically be attracted to our biomaterial and start to conduct regenerative activities. Indeed this approach finds very much attention nowadays in most tissue engineering studies. For example, El-Ghannan and colleagues [122] have shown that fibronectin preferentially adsorbs on calcium-phosphate-coated bioactive glass instead of untreated bioactive glass and free HAp. The increased fibronectin adsorption promoted osteoblast function on calcium-phosphate-coated bioactive glass. It is now well established that osteoblasts preferentially bind to specific amino-acid sequences like Arginine–Glycine–Aspartic acid (RGD) and heparin sulfate binding regions adsorbed proteins. The interactions of osteoblasts with biomaterial surfaces have been reviewed in detail by Anselme [123]. The simplest way to produce nanocomposites is to blend a nano-scale material with another nano- or micro-scale material; however, it is difficult to control homogeneity of dispersion of one phase in the other, which will impart unpredictable effects on the physical or chemical properties of the final composite. Calcium phosphate crystals prepared for use as nanocomposite components are presented in Fig. 4. A number of researchers have tried HAp and collagen composites by anchor- 1264 V. Hasirci et al. Figure 4. Calcium phosphate crystals produced at METU. ing HAp particles in a collagen matrix in order to improve mechanical properties and bioactivity [124 –126]. However, such techniques did not result in a structure similar to natural bone. Bone is composed of nano-sized HAp crystals and collagen fibers, in which the c-axes of the HAp are regularly aligned along the collagen fibers. This nanostructure plays a crucial role in bone metabolism and mechanical properties. Detailed examination of the mineralization process of collagen and having known the conditions for some simple physiological reactions has led scientists to the idea to mimic this natural mineralization process synthetically. Kikuchi et al. [127] have synthesized a composite of HAp with collagen by a simultaneous titration/coprecipitation method using Ca(OH)2 , H3 PO4 and porcine atelocollagen as starting materials. They obtained HAp/collagen composites with a nanostructure similar to bone in which the c-axes of blade-shaped HAp nanocrystals, 50–100 nm in size, were aligned along collagen fibers up to 20 µm in length. This alignment was selfassembled by the chemical interaction between HAp and collagen. The composite had about 40 MPa bending strength and a Young’s modulus of 2.5 GPa, which seems sufficient for bone-graft materials, being the same as autogeneous cancellous bone. The composite was incorporated into the remodeling process of bone in vivo, resorbed by osteoclastic cells, and new bone was formed by osteoblasts after the resorption, as if the composite was grafted autologous bone. Metallic nanocomposites In hard tissue implants, the mechanical properties of biomaterials are the primary important selection criterion just after biocompatibility. They include proper approximation of elastic modulus of implant to bone, sufficient compressive strength and endurance to cyclic loads. Thus, metals have been the first choice for orthopedic implants, since the beginning of orthopedic applications. Although today there are some alternatives, such as polymers or polymer composites, to metals in orthopedic Nanobiomaterials: review of the existing science and technology, and new approaches 1265 implants, their use is limited to non-load-bearing applications at the present time and metals are essential in hard tissue implants. Therefore, currently there are many studies that seek improvements to biocompatibility and tissue response to metals. Loosening of fixations of metal implants is a very important problem in loadbearing applications and tissue in-growth is probably the best possible fixation mechanism compared to widely available solutions like bone cements and mechanical fixations. Cells do not directly adhere to metals and, what is even worse, material debris forming as a result of shear forces around implant surfaces causes tissue necrosis around implant neighborhood. To provide good tissue response, researchers are searching for ways to make composites of metals with bioactive and osteoconductive materials. Recently, He et al. successfully fabricated a titanium composite coated with nano-scale CaP (aggregate)/Al2 O3 for implant applications using a hybrid technique of anodization and hydrothermal treatment [128]. The nanometer scale of the coat was designed both for good biocompatibility and tissue in-growth to achieve an enhanced fixation of implant to host bony tissue. CONCLUSIONS Nanobiomaterials have as diverse applications as we have tried to present above. With the development of new production technologies, application of known methods for creation of nano-products and influx of new knowledge from different areas of science, new uses are being continuously devised leading to better biomaterials and implants and, therefore, higher quality of life. The next decade is expected to be a very exciting period for nano-scale research applications. REFERENCES 1. J. Y. Wong, J. B. Leach and X. Q. Brown, Surface Sci. 570, 119 (2004). 2. E. J. P. Jansen, R. E. J. Sladek, H. Bahar, A. Yaffe, M. J. Gijbels, R. Kuijer, S. K. Bulstra, N. A. Guldemond, I. Binderman and L. H. Koole, Biomaterials 26, 4423 (2005). 3. P. Clark, P. Connolly, A. S. G. Curtis, J. A. T. Dow and C. D. W. Wilkinson, Development 99, 439 (1987). 4. X. F. Walboomers and J. A. Jansen, Odontology 89, 2 (2001). 5. K. Matsuzaka, X. F. Walboomers, J. E. de Ruijter and J. A. Jansen, Biomaterials 20, 1293 (1999). 6. P. X. Ma, Mater. Today 7, 30 (2004). 7. T. Møller-Pedersen, Exp. Eye Res. 78, 553 (2004). 8. J. S. Belkas, M. S. Shoichet and R. Midha, Oper. Tech. Orthoped. 14, 190 (2004). 9. D. L. Wilson, R. Martin, S. Hong, M. Cronin-Golomb, C. A. Mirkin and D. L. Kaplan, Proc. Natl. Acad. Sci. USA 98, 13660 (2001). 10. H. Schift, L. J. Heyderman, C. Padeste and J. Gobrecht, Microelectron. Eng. 61-62, 423 (2002). 11. H. G. Craighead, C. D. James and A. M. P. Turner, Curr. Opin. Solid State Mater. Sci. 5, 177 (2001). 12. D. G. Choi, J. Jeong, Y. Sim, E. Lee, W. S. Kim and B. S. Bae, Langmuir 21, 9390 (2005). 13. Y. N. Xia, J. A. Rogers, K. E. Paul and G. M. Whitesides, Chem. Rev. 99, 1823 (1999). 1266 V. Hasirci et al. 14. A. I. Teixeira, G. A. Abrams, C. J. Murphy and P. F. Nealeya, J. Vac. Sci. Technol. B 21, 683 (2003). 15. E. T. den Braber, J. E. de Ruijter, L. A. Ginsel, A. F. von Recum and J. A. Jansen, J. Biomed. Mater. Res. 40, 291 (1998). 16. B. Chehroudi, D. McDonnel and D. M. Brunette, J. Biomed. Mater. Res. 34, 279 (1997). 17. H. Kenar, G. Torun Köse and V. Hasirci, Biomaterials 27, 885 (2006). 18. A. I. Teixeira, G. A. Abrams, P. J. Bertics, C. J. Murphy and P. F. Nealey, J. Cell Sci. 116, 1881 (2003). 19. M. E. Manwaring, J. F. Walsh and P. A. Tresco, Biomaterials 25, 3631 (2004). 20. M. J. Dalby, M. O. Riehle, D. S. Sutherland, H. Agheli and A. S. G. Curtis, J. Biomed. Mater. Res. 69A, 314 (2004). 21. C. W. Wilkinson, M. Riehle, M. Wood, J. Gallagher and A. S. G. Curtis, Mater. Sci. Eng., C 19, 263 (2002). 22. M. J. Dalby, D. Pasqui and S. Affrossman, IEEE Proc. Nanobiotechnol. 151, 53 (2004). 23. Y. Wan, Y. Wang, Z. Liu, X. Qu, B. Han, J. Bei and S. Wang, Biomaterials 26, 4453 (2005). 24. X. Zong, H. Bien, C. Y. Chung, L. Yin, D. Fang, B. S. Hsiao, B. Chu and E. Entcheva, Biomaterials 26, 5330 (2005). 25. M. J. Dalby, M. O. Riehle, H. Johnstone, S. Affrossman and A. S. G. Curtis, Cell Biol. Int. 28, 229 (2004). 26. M. J. Dalby, N. Gadegaard, M. O. Riehle, C. D. W. Wilkinson and A. S. G. Curtis, Int. J. Biochem. Cell Biol. 36, 2005 (2004). 27. M. Lee, J. C. Y. Dunn and B. M. Wu, Biomaterials 26, 4281 (2005). 28. S. Ber, G. Torun Köse and V. Hasırcı, Biomaterials 26, 1977 (2005). 29. Z. M. Huang, Y. Z. Zhang, M. Kotaki and S. Ramakrishna, Composites Sci. Tech. 63, 2223 (2003). 30. Y. Zhang, C. T. Lim, S. Ramakrishna and Z. M. Huang, J. Mater. Sci. Mater. Med. 16, 933 (2005). 31. R. Dersch, M. Steinhart, U. Boudriot, A. Greiner and J. H. Wendorff, Polym. Adv. Technol. 16, 276 (2005). 32. C. Y. Xu, R. Inai, M. Kotaki and S. Ramakrishna, Biomaterials 25, 877 (2004). 33. M. Li, M. J. Mondrinos, M. R. Gandhi, F. K. Ko, A. S. Weiss and P. I. Lelkes, Biomaterials 26, 5999 (2005). 34. H. Yshimoto, Y. M. Shin, H. Terai and J. P. Vacanti, Biomaterials 24, 2077 (2003). 35. S. R. Bhattarai, N. Bhattarai, H. K. Yi, P. H. Hwang, D. I. Cha and H. Y. Kim, Biomaterials 25, 2595 (2004). 36. S. A. Riboldi, M. Sampaolesi, P. Neunenschwander, G. Cossu and S. Mantero, Biomaterials 26, 4606 (2005). 37. N. Bhatarai, D. Edmonson, O. Veiseh, F. A. Matsen and M. Zhang, Biomaterials 26, 6176 (2005). 38. E. Genove, C. Shen, S. Zhang and C. E. Semino, Biomaterials 26, 3341 (2005). 39. ESA (Electrostatics Society of America), Newsletter 166 (2003). 40. K. S. Rho, L. Jeong, G. Lee, B. M. Seo, Y. J. Park, S. D. Hong, S. Roh, J. J. Cho, W. H. Park and B. M. Min, Biomaterials 27, 1452 (2006). 41. G. E. Martin, I. D. Cockshott and F. J. T. Fildes, US Patent No. 4,044,404 (1977). 42. L. Pinchuk, J. B. Martin Jr. and A. A. Maurin, US Patent No. 5,376,117 (1994). 43. X. Zong, K. Kim, D. Fang, S. Ran, B. S. Hsiao and B. Chu, Polymer 43, 4403 (2002). 44. C. J. Buchko, K. M. Kozlo and D. C. Martin, Biomaterials 22, 1289 (2001). 45. K. Kim, Y. K. Luu, C. Chang, D. Fang, B. S. Hsiao, B. Chu and M. Hadjiargyrou, J. Control. Rel. 98, 47 (2004). 46. G. Verreck, I. Chun, J. Rosenblatt, J. Peeters, A. V. Dijck, J. Mensch, M. Noppe and M. E. Brewster, J. Control. Rel. 92, 349 (2003). Nanobiomaterials: review of the existing science and technology, and new approaches 1267 47. J. Sandler, P. Werner, M. S. P. Shaffer, V. Demchuk, V. Altstadt and A. H. Windle, Composites: Part A (Appl. Sci. Eng.) 33, 1033 (2002). 48. H. Fong, Polymer 45, 2427 (2004). 49. R. L. Price, M. C. Waid, K. M. Haberstroh and T. J. Webster, Biomaterials 24, 1877 (2003). 50. Y. C. Ahn, S. K. Park, G. T. Kim, Y. J. Hwang, C. G. Lee, H. S. Shin and J. K. Lee, Curr. Appl. Phys. 6, 1030 (2006). 51. K. Graham, M. Ouyang, T. Raether, T. Grafe, B. McDonald and P. Knauf, in: 15th Annual Tech. Conference & Expo of the American Filtration & Separations Society, Galveston, TX, p. 1 (2002). 52. Z. Maa, M. Kotaki and S. Ramakrishna, J. Membr. Sci. 265, 115 (2005). 53. K. Sawicka, P. Gouma and S. Simon, Sensors Actuat. B 108, 585 (2005). 54. Y. Xian, Y. Hu, F. Liu, Y. Xian, H. Wang and L. Jin, Biosensors Bioelectr. (2005) (in press). 55. K. K. Jain, Drug Discovery Today 10, 21 (2005). 56. P. Pathak, M. J. Meziani, T. Desai and Y. P. Sun, J. Am. Chem. Soc. 126, 10842 (2004). 57. J. Whelan, Drug Discovery Today 6, 1183 (2001). 58. M. H. Sheridan, L. D. Shea, M. C. Peters and D. J. Mooney, J. Control. Rel. 64, 91 (2000). 59. Y. H. Yuna, D. J. Goetzb, P. Yellena and W. Chena, Biomaterials 25, 147 (2004). 60. R. Kumar, U. Bakowsky and C. M. Lehr, Biomaterials 25, 1771 (2004). 61. E. K. Park, S. B. Lee and Y. M. Lee, Biomaterials 26, 1053 (2005). 62. J. C. Olivier, J. Am. Soc. Exp. Neurother. 2, 108 (2005). 63. G. A. Hughes, Nanomed. Nanotechnol. Biol. Med. 1, 22 (2005). 64. H. M. Courrier, N. Butz and T. F. Vandamme, Crit. Rev. Ther. Drug Carrier Syst. 19, 425 (2002). 65. M. P. Desai, Pharm. Res. 13, 1838 (1996). 66. I. Brigger, C. Dubernet and P. Couvreur, Adv. Drug Delivery Rev. 54, 631 (2002). 67. S. Zhang, Biotechnol. Adv. 20, 321 (2002). 68. H. G. Hansma, K. Kasuya and E. Oroudjev, Curr. Opin. Struct. Biol. 14, 380 (2004). 69. D. Cui, Biotechnol. Prog. 19, 683 (2003). 70. C. A. Mirkin, R. L. Letsinger, R. C. Mucic and J. J. Storhoff, Nature 382, 607 (1996). 71. S. K. Sahoo and V. Labhasetwar, Drug Discovery Today 8, 1112 (2003). 72. A. Lamprecht, N. Ubrich, H. Yamamoto, U. Schäfer, H. Takeuchi, P. Maincent, Y. Kawashima and C. M. Lehr, J. Pharmacol. Exp. Ther. 299, 775 (2001). 73. W. L. Monsky, D. Fukumura, T. Gohongi, M. Ancukiewcz, H. A. Weich, V. P. Torchilin, F. Yuan and R. K. Jain, Cancer Res. 59, 4129 (1999). 74. T. K. Jain, I. Roy, T. K. De and A. Maitra, J. Am. Chem. Soc. 120, 11092 (1998). 75. M. Lal, L. Levy, K. S. Kim, G. S. He, X. Wang, Y. H. Min, S. Pakatchi and P. N. Prasad, Chem. Mater. 19, 2632 (2000). 76. M. C. Jones and J. C. Leroux, Eur. J. Pharm. Biopharm. 48, 101 (1999). 77. J. Wang, D. Mongayt and V. P. Torchilin, J. Drug Target. 13, 73 (2005). 78. N. Nishiyama and K. Kataoka, Adv. Exp. Med. Biol. 519, 155 (2003). 79. K. Kataoka, A. Harada and Y. Nagasaki, Adv. Drug Deliv. Rev. 47, 113 (2001). 80. C. H. Wang, C. H. Wang and G. H. Hsiue, J. Control. Rel. 108, 140 (2005). 81. D. A. Tomalia, H. Baker, J. Dewald, M. Hall and G. Kallos, Polym. J. 17, 117 (1985). 82. A. Quintana, E. Raczka, L. Piehler, I. Lee, A. Myc, I. Majoros, A. K. Patri, T. Thomas and J. Mulé, J. R. Baker, Pharm. Res. 19, 1310 (2002). 83. A. K. Patri, J. F. Kukowska-Latallo and J. R. Baker, Adv. Drug Deliv. Rev. 57, 2203 (2005). 84. R. Duncan and L. Izzo, Adv. Drug Deliv. Rev. 57, 2215 (2005). 85. J. Villen, E. Oliveira, J. I. Nunez, N. Molina, F. Sobrino and D. Andreu, Vaccine 22, 3523 (2004). 86. Y. Hu, Y. Chen, Q. Chen, L. Zhang, X. Jiang and C. Yang, Polymer 46, 12703 (2005). 1268 V. Hasirci et al. 87. J. P. Salvage, S. F. Rose, G. J. Phillips, G. W. Hanlon, A. W. Lloyd, I. Y. Ma, S. P. Armes, N. C. Billingham and A. L. Lewis, J. Control. Rel. 104, 259 (2005). 88. C. Lo, K. Lin and G. Hsiue, J. Control. Rel. 104, 477 (2005). 89. U. O. Häfeli, Int. J. Pharm. 277, 19 (2004). 90. E. S. Gil and S. M. Hudson, Progr. Polym. Sci. 29, 1173 (2004). 91. I. Y. Galaev and B. Mattiasson, Trends Biotechnol. 17, 335 (1999). 92. B. Jeong and A. Gutowska, Trends Biotechnol. 20, 305 (2002). 93. Y. Shin, J. Chang, J. Liu, R. Williford, Y. Shin and G. J. Exarhos, J. Control. Rel. 73, 1 (2001). 94. C. Chaw, K. Chooi, X. Liu, C. Tan, L. Wang and Y. Yang, Biomaterials 25, 4297 (2004). 95. I. Kim, Y. Jeong, C. Cho and S. Kim, Int. J. Pharm. 211, 1 (2000). 96. H. Wei, X. Zhang, Y. Zhou, S. Cheng and R. Zhuo, Biomaterials 27, 2028 (2006). 97. S. Kim, J. Ha and Y. M. Lee, J. Control. Rel. 65, 345 (2000). 98. J. E. Chung, M. Yokoyama and T. Okano, J. Control. Rel. 65, 93 (2000). 99. J. Wu, S. Liu, P. W. Heng and Y. Yang, J. Control. Rel. 102, 361 (2005). 100. H. Vihola, A. Laukkanen, J. Hirvonen and H. Tenhu, Eur. J. Pharm. Sci. 16, 69 (2002). 101. G. Hsiue, S. Hsu, C. Yang, S. Lee and I. Yang, Biomaterials 23, 457 (2000). 102. A. Chilkoti, M. R. Dreher and D. E. Meyer, Adv. Drug Del. Rev. 54, 1093 (2000). 103. P. Gupta, K. Vermani and S. Garg, Drug Discovery Today 7, 569 (2002). 104. M. E. Eccleston, M. Kuiper, F. M. Gilchrist and N. K. H. Slater, J. Control. Rel. 69, 297 (2000). 105. R. A. Siegel, Adv. Polym. Sci. 109, 233 (1993). 106. O. E. Philippova, D. Hourdet, R. Audebert and A. R. Khokhlov, Macromolecules 30, 8278 (1997). 107. J. Dai, T. Nagai, X. Wang, T. Zhang, M. Meng and Q. Zhang, Int. J. Pharm. 280, 229 (2004). 108. A. C. Foss, T. Goto, M. Morishita and N. A. Peppas, Eur. J. Pharm. Biopharm. 57, 163 (2004). 109. N. Murthy, J. Campbell, N. Fausto, A. S. Hoffman and P. S. Stayton, J. Control. Rel. 89, 365 (2003). 110. T. R. Kyriakides, C.Y. Cheung, N. Murthy, P. Bornstein, P. S. Stayton and A. S. Hoffman, J. Control. Rel. 78, 295 (2002). 111. R. Langer and D. A. Tirrell, Nature 428, 487 (2004). 112. M. Stubbs, P. M. J. McSheehy, J. R. Griffiths and C. L. Bashford, Mol. Med. Today 6, 15 (2000). 113. K. Na, K. H. Lee and Y. H. Bae, J. Control. Rel. 97, 513 (2004). 114. K. Na, E. S. Lee and Y. H. Bae, J. Control. Rel. 87, 3 (2003). 115. H. Wakamatsu, K. Yamamoto, A. Nakao and T. Aoyagi, J. Magnet. Magnet. Mater. 302, 327 (2006). 116. T. J. Webster, Am. Ceram. Soc. Bull. 82, 1 (2003). 117. D. F. G. Emery, H. J. Clarke and M. L. Glover, J. Bone Joint Surg. 79, 240 (1997). 118. T. Kokubo, H. M. Kim and M. Kawashita, Biomaterials 24, 2161 (2003). 119. H. Gao, B. Ji, I. L. Jager, E. Arzt and P. Fratzl, Proc. Natl. Acad. Sci. USA 100, 5597 (2003). 120. T. J. Webster, C. Ergun, R. H. Doremus, R. W. Siegel and R. Bizios, J. Biomed. Mater. Res. 51, 475 (2000). 121. T. J. Webster, R. W. Siegel and R. Bizios, Biomaterials 20, 1221 (1999). 122. A. El-Ghannan, P. Ducheyne and I. M. Shapiro, J. Orthoped. Res. 17, 340 (1999). 123. K. Anselme, Biomaterials 21, 667 (2000). 124. C. Du, F. Z. Cui, Q. L. Feng, X. D. Zhu and K. de Groot, J. Biomed. Mater. Res. 42, 540 (1998). 125. W. R. Walsh, J. Harrison, A. Loefler, T. Martin, D. Van Sickle, M. K. Brown and D. H. Sonnabend, Clin. Orthoped. 375, 258 (2000). 126. Q. Q. Zhang, L. Ren, C. Wang, L. R. Liu, X. J. Wen, Y. H. Liu, X. D. Zhang, Artif. Cells Blood Substit. Immobil. Biotechnol. 24, 693 (1996). 127. M. Kikuchi, S. Itoh, S. Ichinose, K. Shinomiya and J. Tanaka, Biomaterials 22, 1705 (2001). 128. L. P. He, Y. W. Mai and Z. Z. Chen, Mater. Sci. Eng. A 367, 51 (2004).