Keywords

2.1 Introduction

Nanocomposites can be defined as an assortment of two or more materials where at least one of those materials should be on a nanometer-scale. By utilizing the composite approach and secondary substitution phases, it is possible to produce nanocomposites with mechanical properties such as Young’s Modulus and strength similar to those of human cortical and cancellous bone. Nanocomposites can be manufactured by either mixing physically or through the incorporation of a new element into an existing nanosized material. This introduction permits the modification of the nanomaterial’s properties and this may in turn offer new function for the material [1,2,3,4].

An alternative form of nanocomposite known as the gel system has been developed and ideal for biomedical applications. It is a three-dimensional network immersed in a fluid and the system can entrap nanostructured materials. Matching the specific requirements of biomedical devices is made possible through the gel approach as the properties of nanomaterials can be enhanced and modified. Additionally, acrylamide hydrogel has been used to entrap indicator dyes for the development of intracellular biosensors [5, 6], while carbon nanotube aqueous gel was developed for applications in enzyme-based biosensors [7].

Nanogel, which is a flexible, nanosized hydrophilic polymer gel, is an example of a gel system that can be applied as drug delivery carriers [8, 9]. According to a study by Vinogradov et al., drugs can be loaded spontaneously into the nanogel after they were synthesized and swollen in water [8]. The creation of dense nanoparticles occurs once the gel collapses due to a decrease in the solvent volume.

For applications in dentistry and tissue engineering, the most widely used materials now are those selected from a dozen or so well-characterized and available biocompatible polymers, metals, ceramics, in addition to their combinations as composites and hybrids. The research and development of new nanocomposites and nanolaminates applicable to implant dentistry and regenerative medicine are driven at an astonishing rate by the combined efforts of novel synthesis approach and the advancement into new enabling methodologies such as surface modification, bioinspired (biomimetics), microscale, and nanoscale fabrications. The last part of the two-part chapter aims to give an overview into the production and types of bio-nanocomposites being examined for applications in implant dentistry as a surface modification approach to improve bone-implant bonding and as tissue engineering scaffolds in the treatment of bone defects. Lastly, brief insights into the application of marine structures in bone grafts and in the treatment of bone disease are also provided.

2.2 Osseointegration of Implants: Issues and Concerns

One of the primary concerns in biomedical materials research is the relationship between surface properties and the biological responses of materials. Currently, dental and orthopedic implants are produced from metals such as titanium alloy and bioinert cobalt chromium alloys. Their major disadvantage is failure to adapt to the local tissue environment and do not chemically bond to bone unless modified. At present, two different methods are used to insert orthopedic implants surgically. The first uses bone cement (predominately poly(methyl methacrylate) or PMMA) for strong adhesion. The second approach utilizes bioactive ceramic coating to coat porous or micro-textured implants for chemical bonding and mechanical interlocking. This technique is extensively applied to dental and maxillofacial implants in addition to orthopedic prostheses.

From a dental perspective, establishing the success criteria for implant systems are vital and to test implants in well-controlled clinical manner. The most obvious sign of implant failure is mobility and its influence on the surrounding bone tissue. In addition, preventing inflammatory responses and osteolysis is a critical requirement for any implant regardless of the material used in its manufacturing (i.e., ceramic, polymer, of metal). Bone remodeling occurs during the first year of function in response to occlusal forces and to establish normal dimensions of the peri-implant soft tissues.

The stability of the implant at the time of insertion is extremely important and after an implantation, a number of tissue responses can take place at the interface between hard or soft tissues and the implant. Numerous factors can affect the success rate of a dental implant or prosthesis such as the properties and structures of the materials used, the surgical procedure or technique employed, the design of the implant or device, biomechanics and application and impact of the functional loading and finally, the medical condition and health of the patient.

2.2.1 Biological Activities and Cellular Responses

While metals such as titanium and its ternary alloys (for example Ti-6Al-4V) have been applied successfully for more than half a century, its relationship with aseptic inflammation particularly in orthopedic joint replacements believed to be the result of titanium particles releasing from the surfaces of implants and prostheses into the surrounding microenvironment has been a concern [10]. It has been revealed that particles could be released in a surface type-dependent fashion after ultrasonic scaling of titanium implants that may aggravate peri-implantitis [10]. Surface modification and the utilization of bioceramic coatings (both nanocoatings and nanocomposite coatings) on these materials are intended to offer protection against the release of metal ions which might trigger a negative host response. Ultimately, this provides an improved environment and structure for the growth of new bone [1,2,3,4, 11, 12].

Concerns and problems during the past three decades related to biological interactions and adhesion to improve the reliability and longevity of implants and prostheses have driven the assessment of surface modification towards bone-implant adaptability, rapid healing, and early osseointegration. Biological fixation is used to define the way the implant or prostheses is bonded firmly to the host tissue through bone in-growth (with or without the help of mechanical fixation) and not requiring the use of any adhesives. Surface improvements by increasing bioactivity through biological and chemical processes and surface micro- and macro-texturing have been the primary emphasis for researchers in the dental and surgical arena.

Studies based on human trials and animal models have demonstrated that the deposition of a thin hydroxyapatite and other calcium phosphate coatings on surfaces of implants accelerated early bone formation as well as an increase in bond strength between bone and implant [13,14,15,16,17,18,19,20,21]. It has been shown that the interfacial area or the healing zone between the hydroxyapatite coating and mature bone was found to be rich in mesenchymal stem cells (MSCs) and collagen after 2 weeks post-implantation, which differentiated into osteoblast phenotype and formed osteoid, the non-mineralized bone indicating the commencement of bone regeneration (Fig. 2.1) [14]. Similar observations were made 4 weeks post-implantation where implants coated with nanostructured calcium phosphate inserted into the rabbit femur generated significantly great bone-to-implant contact relative to non-coated titanium implants [15].

Fig. 2.1
figure 1

Reprinted with permission from [14]

Optical photomicrograph of a longitudinal section of a uncoated titanium and b hydroxyapatite-coated titanium, implanted in rat femur at 2 weeks showing (O) non-mineralized osteoid, and c collagen matrix with Putative Mesenchymal Cells (Goldner’s Masson Trichrome stain).

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Even though many biomaterials have been introduced in dentistry, tissue engineering, and orthopedics, calcium phosphate owing to its similarity to human bone and above all, its dissolution characteristics that enable bone growth and regeneration, holds a special consideration. It should be mentioned that in this chapter the chemical properties would be viewed from the standpoint that hydroxyapatite is calcium phosphate given that most of the information published on hydroxyapatite is categorized under calcium phosphate where hydroxyapatite also belongs, even though hydroxyapatite will have reactivities and properties within the physiological environment different from those of other phosphates.

2.2.2 Surface Modification of Implants and Prostheses

The ability to generate an ideal environment for bone growth has been used to determine if an implant or prosthesis coated with a biocompatible material is successful from a clinical perspective, which involves reducing the amount of metal ions being released, the formation of adequate mechanical interlocking, and the accessibility of a bioactive surface for biological and chemical bonding [1,2,3,4, 11, 12].

As explained in the previous section, calcium phosphate has been demonstrated to be an ideal material for dental and orthopedic coatings in a number of studies and extensive bone apposition in animal models when synthesized in a correct manner, and based on these findings [13,14,15,16,17,18,19,20,21], it can be said that the long-term performance and quality of an implant and/or prosthesis coated with calcium phosphate are determined by various factors such as crystallinity, surface chemistry, constituent phases, surface topography, porosity, and thickness of the coating (Fig. 2.2) [1,2,3,4, 11, 12].

Fig. 2.2
figure 2

Calcium phosphate-coated dental implants

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Although chemically similar to the mineral component of hard tissues and bones in mammals, porous calcium phosphate has unfavorable mechanical properties and consequently, they cannot be utilized in load-bearing applications. Biological apatites are the inorganic phases of calcified tissues such as bones and teeth and have been idealized as calcium hydroxyapatite. Nevertheless, their lattice parameters are different in addition to their association with other ions such as magnesium and CO3 and typically calcium-deficient [11, 12]. This created a motivation for the synthesis and deposition of both thin and thick coatings on metallic alloys. Utilizing calcium phosphate due to its bioactivity as coatings on metallic alloys such as Ti-6Al-4V allows the union of the two materials’ key properties to create a single and functional component [11, 12]. Consequently, it is important to gain an understanding into the mechanisms behind biofixation that takes place when calcium phosphate coating is used. It has been proposed that an increase in the concentration of calcium and phosphate due to the partial dissolution of the apatite into the microenvironment, followed by the formation of carbonate apatite microcrystals and their amalgamation with the organic matrix of bone causing biological growth of bone tissue [22].

2.2.3 Nanocomposite Coatings: Production Techniques

Four general conventional industrial coating techniques were envisioned during the past three decades for the synthesis of bioactive calcium phosphate coatings for clinical applications [11, 12, 23,24,25]. Currently, coating methods such as sol–gel and thermal or plasma spraying have been used in the deposition of calcium phosphate coating, with plasma spraying being the primary deposition technique utilized for medical applications. Despite its extensive use, there were serious concerns around the coatings deposited by plasma spraying, such as the coatings produced are relatively thick, highly porous and contains amorphous phases. More importantly, the deposited coatings are typically non-uniform and bonds poorly to metal implants. Given the fact that high temperature is used during plasma spraying, it is well-known that hydroxyapatite will undergo dissolution to calcium oxide and β-tri-calcium phosphate, which will cause complications within the physiological environment as these phases have much quicker rates of dissolution.

Another concern is a decrease in mechanical properties of titanium substrates when thick coatings are produced as it is essential to sinter powder ceramics at temperatures of 1000 °C or above. If the temperature is below 882.5 °C or the beta phase transus temperature, commercially pure titanium retains a hexagonal close-packed (HCP) crystal structure (alpha phase). On the other hand, titanium will undergo transformation to a body-centered cubic structure (beta phase) if the temperature is higher than the transus temperature. This transformation will create strains within the titanium substrate that degrades the bond strength of the ceramic coating [26]. The newer sol–gel coating process uses much lower temperatures and thus averting the problems caused by the phase transformation of titanium.

To achieve ideal bioactivity and biocompatibility during the production of implant materials is vital and nanocoatings (particularly those synthesized using the sol–gel approach) can offer an efficient and cost-effective approach in altering the interactions of the implant with the “host” environment. Superior strength, hardness, and bioactivity are known to be displayed by nanocrystalline coatings because of the grain sizes falling in the nanometer range [11, 12, 27]. During the early 1990s, several nanoscale calcium phosphate coatings synthesized using the sol–gel technique were introduced ranging from mixed calcium phosphates to 100% pure hydroxyapatite with a mixture of amorphous to finely crystalline phases [27]. The thickness of the coating produced also varied based factors such as chemistry, viscosity, and the coating methods used (Fig. 2.3) [27].

Fig. 2.3
figure 3

Reprinted with permission from [27]

Sol–gel nanocoating of hydroxyapatite on Ti-6Al-4V substrates. The thickness of the coating is approximately 70 nm. a Results of the adhesion testing revealed the behavior of the nanocoating under tensile loads, mainly crack arrest and crack branching, which is indicative of excellent fracture toughness behavior. Please note the slip lines of the titanium alloy and the formation of the cracks in these lines. b Catastrophic failure of the coating and fracture of the titanium substrate under the coating.

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The fabrication of new nanocomposites and nanolaminates that are applicable to an extensive range of applications such as surface-modified dental and orthopedic implants/prostheses for improved soft and hard tissue attachment and scaffolding material with increased bioactivity for tissue regeneration and engineering has been ongoing. Since 2000, the development of nanoceramic composite based on calcium phosphate has been the focus for biomedical and dental researchers. Typically, the thickness of a single-layered coating is less than 100 nm. Coatings that contain multiple layers with suitable mechanical, biological, physical, and chemical properties can be manufactured with relative ease and can be applied with different compositions as multi-layered gradient coatings or nanolaminates [1]. As such, one form of nanocomposite coating is a multilayered, nanolaminated mixed nanocoatings and this can be manufactured by laminating various nanocoatings together to achieve the desired properties, structures, and thickness [1].

The properties of the calcium phosphate-based nanocomposite coatings and nanolaminates are dependent on the dispersion or secondary phase. Currently, the development of new generations of nanocomposite coatings containing synthetic and natural nanomaterials such as bioglass, collagen, and chitosan are being pursued to promote osseointegration [4]. Polymeric materials have also been applied as the dispersion phase in nanocomposites with calcium phosphate. A study has hypothesized that cellular responses at the bone-titanium implant interface can be induced if a maleic polyelectrolyte such as sodium maleate-vinyl acetate copolymer is added to calcium phosphate. Cells grown on the coated surfaces induced a higher proliferation rate and the incorporation of a sodium maleate copolymer improved surface bio-adhesion than compared to just calcium phosphate [28].

2.2.3.1 Biological Materials

A greater emphasis will be focused on the addition of molecular and nanoscale-based biological materials such as bone morphogenetic proteins, liposomes, and peptides to calcium phosphate coatings in the future in an attempt to reduce the timeframe for implant integration and to improve and promote osseointegration of dental and orthopedic implants and prostheses. Moreover, the current trends are centered on the incorporation of stem cells, growth factors, bone morphogenetic proteins, and several pharmaceuticals into multifunctional nanocoatings.

2.2.3.1.1 Chitosan

The bioactivity of chitosan, despite the fact that it is one of the most widely examined material in tissue engineering, will require enhancements for certain tissues and therefore making them not as ideal if used as a stand-alone material [29]. More importantly, the poor mechanical properties of chitosan have restricted their use in load-bearing applications and its combination with reinforcing bioactive materials is designed to reduce these weaknesses. For that reason, it is added to calcium phosphate to overcome these shortcomings [4]. Studies have shown that composite coatings of calcium phosphate and chitosan demonstrated increased osteoconductivity and biodegradation along with sufficient mechanical strength [30, 31].

2.2.3.1.2 Collagen

It has been suggested that the unique nanocomposite structure of human bone tissue could be replicated by a composite consists of calcium phosphate and collagen, and this combination can be beneficial as the ductile properties of collagen counterbalance the poor fracture toughness of calcium phosphate [32,33,34,35,36,37,38]. In 2010, a study has revealed the osteogenic behavior of rat bone marrow cells could be stimulated in vitro by calcium phosphate-collagen composite coatings even if the coating thickness is less than 100 nm. The composite coating is also able to enhance osteoblast differentiation compared to pure calcium phosphate coatings based on accelerated mineral deposition and a decrease in proliferation [36]. In another animal model study, observations after four weeks post-implantations revealed calcium phosphate-collagen nanocomposite coated titanium rods were virtually surrounded by new bone tissue without encapsulation and had the highest ratio of bone contact when placed under the periosteum of a rat calvarium, while calcium phosphate-coated as well as uncoated specimens were encapsulated with fibrous tissues. Additionally, the nanocomposite-coated rods generated the highest bonding strength to bone [35].

2.2.3.1.3 Bone Morphogenetic Proteins

After sintering at a high temperature, commercial calcium phosphate ceramics display a very low specific surface area as well as poor surface reactivity. Despite calcium phosphate is an exceptional osteoconductive material, their osteoinductive properties seem rather weak. It has been proposed that the biological and surface properties of these ceramics can be improved by coating the surfaces with nanocrystalline carbonated apatite and such a layer could encourage biological activity by stimulating the creation of nanopores and increasing the specific surface area [39]. Enhancements in surface reactivity were verified in an adsorption investigation utilizing an osteogenic growth factor, recombinant human bone morphogenetic protein-2 (rhBMP-2). The study also revealed the ceramic is capable of adsorbing more of the protein and releasing it in a prolonged fashion. It was also concluded that the coated ceramic combined with rhBMP-2 could improve bone formation as observed in an in vivo ovine animal model.

Over the last decade, rhBMP-2 has become one of the most extensively investigated protein for applications in oral and maxillofacial surgery such as maxillary sinus floor augmentation, alveolar grafting, and mandibular reconstruction [40,41,42,43,44,45,46,47,48,49]. Bone morphogenetic proteins (BMPs) are multi-functional growth factors that belong to the superfamily of transforming growth factor β (TGFβ) [50, 51]. In recent years, their roles in cellular functions in postnatal and adult animals have been comprehensively examined [50, 51]. The actions of BMPs were first realized in 1965 [52], but it was not until the late 1980s after the replication of human BMP-2 and BMP-4 in addition to the sequencing and purification of bovine BMP-3 (osteogenin) that the protein accountable for bone induction were identified [53,54,55]. Isolated originally from demineralized bone matrix, rhBMP-2 possesses the capacity to stimulate cell differentiation into chrondroblasts and osteoblasts to begin the process of forming new cartilage and bone tissue [53, 56, 57].

In vivo studies have revealed the application of a calcium phosphate coating (either deposited using conventional coating technique or synthesized biomimetically) can potentially be used to transport growth factors such as BMP in an attempt to enhance peri-implant bone regeneration and their gradual release from the coating has been hypothesized to be the result of cell-mediated degradation [58,59,60,61,62]. Furthermore, BMP can be incorporated and signaling factors be steadily released from the coating during in vivo degradation [60,61,62]. Also, it was postulated that the combination of calcium phosphate coating and rhBMP-2 could result in significant improvement in bone apposition if the implant is immersed into a protein solution before implantation [60]. In vivo observation confirmed coatings containing rhBMP-2 generated the greatest bone-to-implant contact and bone area fraction occupancy after three weeks of implantation. However, the study also revealed there was not a significant enhancement in bone response if rhBMP-2 was adsorbed directly onto the surfaces of titanium implants but still noticeably higher than uncoated implants.

The deposition of a calcium phosphate layer based on precipitation in a simulated body fluid (SBF) solution has also created new opportunities for the incorporation of BMP-2 to enhance the osseointegration of dental implants. Studies were carried out to gain an understanding into the in vivo osteoinductive capacity of these “self-assembled” or biomimetic coatings containing BMP-2 deposited onto titanium implants [61, 62] and more recently on zirconia [58]. Using a well-established ectopic (subcutaneous) ossification model in rats, the biomimetic coating display the capacity to release BMP-2 at a constant rate in vivo and in a sufficient amount to induce bone formation at an ectopic site. Most of all, their findings also suggested that the coating is able to sustain this osteogenic activity for an extended period as well as carrying out this task with a high degree of efficiency and at a low pharmaceutical level. However, their later study revealed there is no substantial influence on the osteoconductivity during the early phases of osseointegration of implants if calcium phosphate is used as a coating to deliver BMP-2 [61]. The bone-interface coverage and the volume of bone deposited within the peri-implant space were greatest for calcium phosphate coated implants that contains no BMP-2. Uncoated implants along with BMP-2 adsorbed calcium phosphate coated implants produced the lowest quantity of bone deposited and bone-interface coverage.

As previously suggested, the combined efforts of osteoconductive apatite and osteoinductive BMP-2 can create a much greater osteogenic surface environment on titanium surfaces [63]. In order to maintain the rate of release of growth factors, BMP-2 is coupled with negatively charged chondroitin sulfate to create a BMP-2 nanocomplex. Titanium coated with calcium phosphate that contains BMP-2 nanocomplex demonstrated faster cell proliferation once mouse osteoblast cells were seeded on the surfaces in comparison to titanium alone and titanium surfaces only coated with calcium phosphate. The gene expressions of bone-specific markers such as type I collagen and osteocalcin were significantly upregulated by the utilization of BMP-2 nanocomplex. Similar observation was also noticed during the examination of ALP activity.

The release of growth factors such as the ostogenic BMP-2 and the angiogenic vascular endothelial growth factor (VEGF) from bioceramics delivery vehicles such as calcium phosphate has been well documented to influence the magnitude of bone formation in animal models [64,65,66,67,68]. Observations from a recent in vivo study has suggested the use of calcium phosphate for the combined delivery of both BMP-2 and VEGF can result in considerable increase in the formation of new bone as well as enhancements in osteogenic potential [64]. However, the effectiveness of calcium phosphate especially in the form of a coating and their utilization to transport growth factors aimed at improving osseointegration at the bone-implant interface remained unclear. Using biomimetically octacalcium phosphate-coated implants, the efficacy of dual delivery of recombinant human BMP-2 and recombinant human VEGF on osseointegration were assessed [67, 68]. After 2 weeks of post-implantation into frontal skulls of domestic pigs, bone volume density values were improved for biomimetically-coated implants containing the combination of recombinant human BMP-2 and recombinant human VEGF but did not significantly improve bone-implant contact after 4 weeks post-implantation [68]. They also suggested based on the results of their later in vivo study that the dual delivery of growth factors favored bone mineralization as well as expression of vital bone matrix proteins [67].

2.2.3.1.4 Peptides

Besides using BMPs to bio-functionalize surfaces of implants, biomimetic peptides such as P-15 and RGD have also been studied for their possible roles in improving the cellular interactions with biomaterials. More importantly, they can be synthetically manufactured, and purification can be carried out relatively easily [69,70,71,72]. Within these biomimetic active peptides, only a cell-binding sequence is contained and by imitating cell-binding sites biologically, the RGD (Arg-Gly-Asp) peptide has been hypothesized to promote cell adhesion [73].

Calcium phosphate coatings combined with an RGD-containing peptide deposited on implant surfaces could potentially enhance the attachment and differentiation of osteoblast. As hypothesized, the osteoconductivity of calcium phosphate was strengthened by the availability of peptide. Once implanted into bone, the protein absorption on calcium phosphate surfaces is altered after the peptide pretreatment [71]. The enhancement in osseointegration may be due to the preferential attachment of cells such as osteoprogenitors to calcium phosphate surfaces. It has also been postulated that the immobilization of RGD on anodized titanium via chemical grafting could enhance the osseointegration of implants [74].

On the contrary, observations from another study revealed there is only weak evidence supporting titanium implants coated with RGD peptides could possibly enhance peri-implant bone formation in the alveolar process despite in vitro studies hypothesizing the use of RGD results in improvements in cell attachment [75]. Similarly, in vivo observations were also noticed as an increase in bone density was not recorded outside the bone-implant interface even though a significant bone stimulating effect can be generated by cyclic RGD at the interface [76]. Furthermore, the potential benefits concerning the use of RGD peptides to functionalize implants coated with calcium phosphate and their effect on osseointegration was questioned [77]. Their results demonstrated the presence of RGD in calcium phosphate disks implanted into rat tibiae significantly inhibited bone formation in addition to the amount of new bone in direct contact with the implant perimeter after 5 days post-implantation [77]. These marginal healing responses has been hypothesized to be due to the uncontrolled signaling responses at the implant-tissue interface by unregulated or sub-optimal integrin binding [78]. In addition, observations from another study proposed structural changes in peptides adsorbed onto titanium as a response to its surface characteristics or low peptide adsorption reduces the likelihood of osteoblast adhesion to be promoted [79].

As a result of the contradictory clinical results produced by the RGD peptides, the search for an alternative peptide sequence that can be applied as an implant coating is a goal worth pursuing. In comparison to RGD-containing peptides, studies have demonstrated that a biomimetic bone matrix known as P-15 was more potent when competing with collagen for cell binding [69, 70, 80]. This biomimetic matrix is a synthetic, 15-amino-acid residue peptide that is identical to the 766GTPGPQGIAGQRGVV780 sequence of the type I collagen α1(I) chain [80]. In 2010, a study was carried out to examine the hypothesis that faster osseointegration process can be achieved if surfaces of dental implants were coated with calcium phosphate and P-15 peptides. After 14 and 30 days of post-implantation into the forehead region of 12 adult pigs, a significant higher percentage of bone-to-implant contact was recorded on implants containing high concentration of P-15 as revealed by the results of histomorphometric analysis. An increase in peri-implant bone density at 30 days was detected on implants containing both high and low concentrations of P-15 peptide [70]. Using a canine model, these observations were later endorsed in a study that confirmed the theory that the bioactivity of implants coated with calcium phosphate can be enhanced by the presence of P-15 and its influence were especially noticeable during the early stages of the healing period [69].

2.2.3.1.5 Stem Cells

The primary challenge related to the clinical application of mesenchymal stem cells (MSCs) is the way they are obtained from healthy tissues. The oral cavity in recent years has played a significant role as a vital source of MSCs. The ease of access to the dental surgeons and clinicians in addition to the efficacy of cells being isolated from dental tissues such as freshly extracted teeth render the clinical utilization of oral-derived stem cells such as dental pulp stem cells, periodontal ligament (PDL) stem cells, and dental follicle progenitor cells extremely attractive [1, 81]. Recently, human periapical inflammatory cysts, which are a biological waste intended to be eradicated surgically to avoid disabling pathological conditions in the oral cavity, also exhibit MSC-like properties. MSCs isolated from human periapical cysts have been shown to differentiate into adiopocytes and osteoblasts based on their self-renewal capacity and multi-lineage differentiation potency [82,83,84]. It has been hypothesized that by combining autologous stem cells from periapical cysts with a bioactive material such as calcium phosphate could provide a solution that encourages the regenerative healing of oral structures such as alveolar bone [82].

Isolated from extracted teeth, PDL contains stem cells which possess the capability to regenerate cementum and/or PDL-like tissues in vivo based on the observations of an animal study [85]. It has been postulated that an alternative implant therapy could be possible through the combined efforts of dental implants with a cell sheet technique [86]. Transplantation of surface-treated and calcium phosphate-coated commercially pure titanium implants with adhered PDL-derived cells (which contain multipotential stem cells) into bone defects in athymic rat femurs as a xenigeneic model as well as into canine mandibular bone as an autologous model were carried out in a previous study to examine the feasibility of regenerating cementum and PDL in vivo. Observations from the rat model demonstrated that PDL-like and cementum-like tissues was partly noticed on surface-treated and calcium phosphate-coated implants combined with adherent PDL-derived cell sheets. Furthermore, histological observations from the canine model revealed the formation of PDL-like and cementum-like tissues was induced on surface-treated and calcium phosphate-coated implants. It was also noticed that the PDL-like tissue was perpendicularly oriented between the titanium surface with cementum-like tissue and the bone [86].

2.3 Regeneration of Hard Tissues: Scaffold Design and Requirements

Tissue engineering in recent times has adopted a new objective by capturing the benefits of uniting the application of three-dimensional bioceramic scaffolds with living cells to transport essential cells to the damaged sites within the patient. During the past few decades, exceptional manufacturing policies have been introduced to amalgamate biogenic materials into bioceramic implants in the areas of osteogenic cell growth and differentiation.

The advancement of bone tissue engineering substitutes has been met by several clinical arguments such as a necessity for improved filler materials used to reconstruct large bone defects and for implants and scaffolds that are more mechanically suited to their biological surroundings. From a biological perspective, bone defect management typically involves the utilization of autograft and allograft [87,88,89]. The purpose of a scaffold is to encourage the formation of new bone from the surrounding tissues as well as providing a template or carrier for implanted bone cells and other biological agents. Although the inclusion of prerequisites regarding the types of materials used during the design of tissue engineering scaffolds are common, it is also vital to include any clinical requirements so that medically relevant bone substitutes can be manufactured. In addition to material and design considerations, bone regeneration requires four biologically vital mechanisms:

  1. 1.

    A morphogenetic signal

  2. 2.

    The willingness of the responsive host cells to react to the signal

  3. 3.

    A suitable carrier that will deliver this signal to specific sites; and

  4. 4.

    Scaffold capable of supporting the growth of responsive host cells and providing a well-vascularized host bed.

The process of bone regeneration is common during fracture repair. It is normal for the evolution of bone regeneration to take place during the repair of fractures or bony defects. The steps involved in the remodeling cycle comprises of the introduction of bone grafts, the skeletal homeostasis, and the cascading cycle of biological events [1]. Sustaining the integrity of the skeleton has been accepted as the principle behind the bone remodeling sequence. This is accomplished through the collaborative efforts of osteoblasts and osteoclasts, and this close collaborative during the remodeling process is frequently referred to as the basic multicellular unit. Bone resorption and bone formation are balanced in a homeostatic equilibrium. The synchronized biological and mechanical actions on a cellular level play a governing role in the delicate equilibrium between bone formation, growth, and resorption. Additionally, the combination of osteoblasts and osteoclasts also contributes to the bone remodeling of defects such as micro-fractures [1].

Derived from cells of embryos, fetuses, or adults, stem cells possess the capacity to replicate for extended periods as well as creating specialized cells that provides the structures for organs and tissues of the human body. Moreover, these stem cells have been proven capable of creating highly vascularized bone tissues when it is combined with mineralized three-dimensional scaffolds and implanted into immune-deficient mice. Treatments of defects across bone diaphysis can be carried out using these bioceramic-cell cultured composites with positive clinical outcomes such as excellent integration of the bioceramic scaffold with bone tissue and good functional recovery. Porous materials such as calcium phosphates are being applied as bone grafts at a rapid rate as they allow the ingrowth of natural bone as well as the supply of nutrients and blood, hence providing a strong bond between bone tissues and the graft material.

Stem cells have been incorporated into a number of bioceramics including calcium phosphate. The development of highly vascularized bone tissues is possible once these bioceramics combined with mineralized three-dimensional scaffolds are implanted into the human body. Such cell cultured-bioceramic nanocomposites can be utilized to treat full-thickness gaps in long bone shafts. More importantly, excellent integration between bone tissue and scaffold can be provided by these bioceramic nanocomposites, and ultimately good functional recovery. The reconstruction of bone tissues utilizing nanocomposite bone grafts with compositional, biological, physiochemical, structural and biomechanical characteristics that mimic those of human bone is an objective to be achieved.

2.3.1 Materials Selection: From Synthetic to Natural

Over the past three decades, a considerable amount of attention has been paid to the development of bioactive composite grafts consist of a polymeric matrix and a bioactive ceramic filler. These bioactive composite grafts are designed to achieve interfacial bonding between the scaffolding graft and the host tissues. Furthermore, such a combination would perfectly bring together the advantages of the two materials in which the polymer would enhance the toughness and the bioceramic would improve the bioactivity of the composite [90].

2.3.1.1 Calcium Phosphate

The repair of bone and periodontal defects, maxillofacial reconstruction, bone space fillers, alveolar ridge augmentation, ear implants, and bone cement additives are just a few examples where calcium phosphate bioceramics are being used in the orthopedic and dental arenas [1]. Calcium phosphates that contain interconnecting pores with diameters between 100 and 500 μm are frequently used as bone graft materials, where they are combined with both natural and synthetic polymers such as chitosan and collagen to form bioactive composite grafts for tissue engineering applications [91, 92]. The dissolution rate of calcium phosphate is determined by its chemistry and structure, this in turn governs the in-situ strength and long-term stability [11, 12, 93].

Composite grafts consist of collagen and calcium phosphate has been of particular interest and a natural option for bone grafting [94]. This is because skeletal bones are composed primarily of carbonate-substituted hydroxyapatite and collagen, both of which are osteoconductive components. Consequently, a scaffold manufactured from calcium phosphate and collagen is anticipated to perform in a similar fashion. Furthermore, collagen-calcium phosphate composite has been shown to be biocompatible in both humans and animal trials. Their study showed calcium phosphate-collagen composites in comparison to monolithic hydroxyapatite displayed osteoconductive properties once embedded with human-like osteoblast cells and generated calcification of an identical bone matrix [95]. Later, another study was carried out to investigate the effect of osteogenic differentiation using bone marrow-derived MSCs on bone regenerative nanocomposites constructed using nano-amorphous calcium phosphate, nano-hydroxyapatite, and reconstituted collagen [96]. It was observed that the presence of reconstituted collagen within the composite significantly improved the osteogenic differentiation.

A number of approaches have been applied to produce gels and films of collagen-calcium phosphate composites as well as collagen-coated calcium phosphate, calcium phosphate-coated collagen matrices and composite scaffolds for bone tissue repair [97]. Based on the observations of several investigations, it has been hypothesized that the utilization of composites composed of nano-hydroxyapatite and tricalcium phosphate with a simple incorporation of collagen could offer improvements in bioactivity as well as the physical properties of the composite scaffold [98,99,100].

First used for bone repair in the early 1900’s with great success, it took tricalcium phosphate more than six decades after the first clinical study that a suggestion was made related to the application of porous calcium phosphate scaffolds to treat bone defects [101, 102]. Presently, commercially available synthetic calcium phosphate biomaterials are classified according to their composition and these include α- and β-tricalcium phosphate, hydroxyapatite, and biphasic calcium phosphate (which is a mixture with variable ratio of hydroxyapatite and β-tricalcium phosphate). Other commercially available calcium phosphate biomaterials have been synthesized from biological and marine materials such as marine algae, hydrothermally converted coral, processed human bone, and bovine bone [11, 12, 93].

In addition to collagen, other biodegradable natural polymers have also been postulated and explored as scaffolding materials in combination with calcium phosphate for bone tissue engineering applications [1, 2]. One such polymer is chitosan and the possibilities of combining nano-calcium phosphate with chitosan to create a composite scaffold for bone regeneration have been extensively investigated [103,104,105]. Synthesized from chitin, a natural polysaccharide that can be isolated from the exoskeletons of crustaceans [106, 107], chitosan contains a structure that is similar to glycosaminoglycan, a vital element of the extracellular matrix and plays a pivotal part in bone regeneration. Observations from a cell culture study have suggested that a calcium phosphate-chitosan scaffold can support the attachment and proliferation of MSCs derived from rat bone marrow and the density of MSC on the scaffold increased by an order of magnitude after 2 weeks [108]. Their study also revealed MSCs attached on the scaffold were able to effectively differentiate down the osteogenic lineage and expressed high levels of ALP. Similarly, findings from different studies also showed MSCs and osteoblast-like cells displayed good proliferation and adhesion onto chitosan-nano-hydroxyapatite composite scaffolds [109,110,111]. In vivo investigations have also been carried out in an effort to study the bone regeneration capabilities of the nanocomposite and observations from the test results revealed that it might not be as effective as using just nano-hydroxyapatite powder [112] or incorporating MSCs to the nanocomposite scaffold prior to implantation [113].

Considered as one of the most abundant renewable resource and the main component in the rigid cells walls in plants, cellulose has been extensively applied in the pharmaceutical and biomedical industries [2]. Observations from in vitro studies have suggested that a nano-calcium phosphate-cellulose composite scaffold is able to support various cellular activities such as growth, attachment, and proliferation of cells such as human dental follicle cells [114] and osteoblasts [115]. A recent in vivo study using rabbit calvarial defect model demonstrated the potential of a cellulose-nano-hydroxyapatite to form mineralized tissues after loaded into defects [116]. Additionally, bacterial cellulose has also received a considerable amount of attention in tissue engineering and their combination with nano-calcium phosphate as a possible composite scaffold for bone regeneration have been examined both in vitro and in vivo [117,118,119].

A well-known fiber used widely in the textile industries, the silk from silkworms has also find applications in tissue engineering due to their excellent biodegradability and biocompatibility. The idea of combining silk fibroin with nano-calcium phosphate was first suggested in 2004 where it was observed that L929 fibroblast cells were able to adhere more abundantly on silk fibroin coated with nano-sintered hydroxyapatite particles [120]. Similar observations were made in another study in which mesenchymal cells adhered and actively proliferated on nano-hydroxyapatite-coated silk fibroin composites [121]. Later, in vitro cell culture study using rat osteoblast cells revealed nonwoven silk fibroin net-nano-hydroxyapatite composite scaffold can enhance the viability of osteoblasts and demonstrated excellent cytocompatibility for cell growth [122]. It has been hypothesized that stem cells could be directed towards osteogenesis using a blended eri-tasar silk fibroin nanofibrous scaffolded with surface precipitated nano-hydroxyapatite. An improved osteogenic differentiation was noticed via ALP assay as well as pattern of gene expression linked to osteogenic differentiation and morphological observations of differentiated cord blood human MSCs under microscopic examination [123]. In addition, in vivo biological performance of silk-nano-calcium phosphate scaffolds was also examined using animal models [124, 125]. Implantation of the scaffold into rabbit knee critical size osteochondral defect models demonstrated a firm integration into the host tissue and de novo bone ingrowths and vessel formation were recorded [125]. Similar bone formation capability was observed when nano-hydroxyapatite mineralized silk fibroin scaffold graft was used to induce greater bone formation and to decrease the height of alveolar bone resorption after tooth extraction in vivo [124].

Apart from natural biodegradable polymers, the manufacture of nanocomposite grafts using synthetic polymers and nano-calcium phosphate have also been postulated and examined [1, 2]. Several advantages are associated with the use of synthetic polymers such as the modification of their physical–chemical properties as well as having the ability to engineer the mechanical and degradation characteristics simply by altering the chemical composition to match specific requirements. Furthermore, their ability to be bioactivated with certain molecules by amalgamation of functional groups and side chains is also another benefit of utilizing synthetic polymers.

The possibility of utilizing high-density polyethylene with calcium phosphate nanoparticles as scaffolds was examine in an in vitro study using human MSCs [126]. The lack of cytotoxic effects of the scaffolds was confirmed by viability assays. After eosin and hematoxylin staining, microscopic images of the cell culture study for six weeks confirmed typical growth and morphology.

The utilization of PLGA or poly(lactic-co-glycolic acid) has sadly been severely limited as a consequence of their poor bioactivity and hydrophobic surface even though it is regarded as one of the most frequently applied biodegradable polymer that can be processed easily. Consequently, a composite has been hypothesized that combines the physiochemical and structural advantages of nano-calcium phosphate in the formation of new bone with the biocompatibility of PLGA. The resultant composite has been believed to be capable of providing a solution in regulating the adhesion and osteogenic differentiation of human MSCs [127,128,129]. Observations from an in vivo study using avian vessels from the chick chorioallantoic membrane have suggested that human adipose-derived stem cells are capable of completely penetrating an amorphous calcium phosphate nanoparticle-PLGA composite within one-week post-implantation [127].

Due to its low inflammatory response, sustained biodegradability, and excellent mechanical characteristics [130], polycaprolactone presents an interesting option for tissue engineering applications [131,132,133]. However, their poor capacity to induce osteogenic cell adhesion, differentiation, and proliferation warrant the creation of a composite that combines polycaprolactone with a bioactive material such as calcium phosphate to overcome these restrictions [133]. It has been shown that the viability and adhesion of human MSCs were enhanced by the presence of nano-hydroxyapatite within the scaffold as revealed by Almar Blue assay as well as higher levels of ALP activity after 14 days of incubation [133]. Moreover, an in vivo study has postulated that the addition of bone marrow-derived MSCs from the patient to a polycaprolactone-nano calcium phosphate scaffold could result in a better environment to promote bone regeneration after implantation into defect site. Observations from a rat critical-sized calvarial defect model revealed noticeably greater bone volume and mineralized regeneration after 2 months post-implantation [131].

2.3.1.2 Bioactive Glass

The work of Hench and his co-workers has led to the discovery of several bioactive glass and glass ceramics, and many clinical Bioglass® and other glass ceramics with similar structures and compositions are being applied in orthopedics as well as in dental and oral and maxillofacial surgery for bone augmentation and restoration. They have also been utilized in the field of tissue engineering in general [134, 135].

Silicon (Si), as one of the important trace elements within the human body, is found at a level of 200–550 ppm attached to extracellular matrix compounds and 100 ppm in the bone [1]. They play a role in the mineralization process of bone growth and they have been discovered in locations of active calcification sites [136]. During the past decade, a new category of bioactive calcium silicate ceramics has been created driven by the bioactive compositions of silicate-based bioglass, the function of silicon within the human body and the behavior of stem cells on silicon mesoporous and nanoporous matrices. Bioactive silicate ceramics with certain compositions has been discovered to promote significantly in vitro osteogenic differentiation of several stem cells and in vivo angiogenesis and osteogenesis. Moreover, studies have been carried out to examine the feasibility of implanting silica-based bioglasses that may also contain small quantities of other crystalline phases [1].

First suggested in 1971, a bioglass known as 45S5 Bioglass® was created (Fig. 2.4) and it was proposed that this glass with a composition of 42% SiO2, 24.5% CaO, 6% P2O5, and 24.5% Na2O by weight has superior osteoblastic activity that allows for a quicker exchange of alkali ions at the surface with hydronium ions in comparison to hydroxyapatite [137, 138]. Furthermore, the formation of an apatite layer was also noticed in glass ceramics with a similar composition and different degrees of crystallinity [139]. Bioactive glasses with reduced alkaline oxide content and precipitated crystalline apatite can be synthesized using a specific heat treatment technique. The glass ceramic produced is called Ceravitals, and it possesses greater mechanical strength, but reduced bioactivity compared to Bioglass®.

Fig. 2.4
figure 4

Modified and adapted from [145]

SEM images of sol–gel-derived agglomerated 45S5 Bioglass® nanoparticles taken at different magnifications.

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Proper glass-to-bone bonding has been demonstrated in bioactive glasses with a certain compositional range containing CaO, SiO2, Na2O, and P2O5 as well as in specific proportions. Results from numerous studies has reached the same conclusion that a firm and unique bonding between the bioactive glass and human bone is achieved as a consequence of the interfacial chemical reactions on the glass surface. This solid and unique bonding was hypothesized to be due to the formation and the even dissolution of a calcium phosphate layer connecting the bioactive glass to the host bone tissue. This theory can be confirmed by examining the bond strength of implants manufactured from both hydroxyapatite and bioactive glass [140,141,142].

An important asset of bioactive glass is its capacity to bond firmly to bone via chemical reaction and this eventually allows the glass to be replaced by advancing bone tissues, making it ideal for applications in medical applications. Of vital significance is the elements found within the glass are minerals or physiological chemicals detected in the human body such as silicon, oxygen, magnesium, sodium, potassium, calcium, and phosphorus. It has been demonstrated that the concentration of those elements never increases to a concentration that could lead to a chemical imbalance in the tissues surrounding the glass implant during bonding and bone formation [143, 144].

The brittle nature of glass subsequently means that it cannot be used in scenarios where load-bearing properties are essential. This led to the development of bioactive glass composites. Nanofibers and nanoparticles of bioactive glass have been made available several years ago and they have been used in conjunction with polymers in the form of a nanocomposite. The rationale behind it is to combine and enhance the bioactive phase of glass particles or fibers with the excellent properties of polymers to improve characteristics such as flexibility and being able to withstand deformation under loading. Nano-bioactive glasses have been manufactured using numerous approaches including the sol–gel approach, gas phase or flame spray synthesis, and laser spinning technique [1, 145].

A number of bioglass-based nanocomposites have been manufactured and examined for applications in orthopedics such as bone regeneration matrix and scaffolds [146,147,148,149,150,151,152] and in dentistry for periodontal and dental-pulp regeneration [153, 154]. Improvements in degradation kinetics, mechanical properties, bioactivity, and osteoblast responses have been observed due to the presence of bioactive glasses in the nanocomposites when compared to pure polymer or bioglass [151, 155,156,157].

2.3.1.3 Marine Materials

Highly functional architectural structures with interconnected open pores can be easily found within the marine environment. These structures are ideal for human implantation in its original form or converted to materials more suitable for biomedical applications due to their high mechanical strength and chemical compositions. Furthermore, off-the-shelf organic and inorganic marine skeletons contains an ideal environment for the proliferation of added MSC populations and promoting bone formation that is appropriate from a clinical perspective. The manufacture of highly efficient scaffolds that can perform at the macro-, micro-, and nanoscopic level will play a vital role in making regenerative medicine a clinical success in the future. Cells will be able to be rearranged and reorganized by these scaffolds into tissues and the encapsulated chemical signals will be released in a targeted fashion and conveying them into the body.

Converted coral skeletons and coralline apatites are perfect examples [158,159,160]. As templates for tissue reconstruction, they have demonstrated substantial clinical success, and this has encouraged tissue-engineering researchers to explore other marine skeletons with enhanced biological and/or mechanical properties. These unique 3-D marine structures can support the growth and enhancement in differentiation of stem cell progenitors into bone cells. This is different to standard carbonate frameworks as they do not induce stem cell differentiation.

Molecules necessary for the regulation and guiding bone morphogenesis and particularly the actions associated with the mineral metabolism and deposition were also discovered in the earliest marine organisms. This is due to the fact that they signify the first molecular components known for calcification, morphogenesis, and wound healing. Bone morphogenetic protein (BMP), the primary cluster of bone growth factors for human bone morphogenesis, are extensively used in musculoskeletal tissue engineering to promote bone tissue formation and gene expression [161]. It has become apparent that BMP are secreted by endodermal cells into the developing skeleton. The use of biochemical factors to trigger cell proliferation and differentiation is one of the most important factors driving the advancement of tissue regeneration.

The design and availability of marine materials have played a critical role in the establishment of one of the simplest solutions to crucial problems in regenerative medicine and in providing frameworks and opportunities of nanofibers, mineralizing proteins, micro- and macrospheres, and osteopromotive analogues. This is demonstrated by the biological efficiency of marine structures such as corals, sponge skeletons, and shells to accommodate self-sustaining musculoskeletal tissues and to promote bone formation through the application of nacre seashells and sponging extracts [158, 160].

The key idea behind biomimetics is to replicate in a laboratory the structures of selected inorganic biomatrices as they play a distinctive role in the fabrication of calcified tissue replacements [162]. Biomaterials in nature contain advantageous properties such as sophistication and complexity, and we are slowly discovering techniques to mimic nature and establish similar levels of complexity although it is to a limited extent. This can be achieved using techniques found in biomineral-inspired materials chemistry. The plan is to use molecules to manufacture skeletons and turning them into macroscopic structures by applying consecutive developmental pathways of systems that nature employs.

Configuring the material environment at the molecular and macro-molecular levels in an attempt to mimic native extracellular matrix is another method that have been studied and the goal is to further expand and translate this continuing research into the creation of clinically relevant scaffolds for regenerative medicine using a unique set of self-organizing hierarchical structures conceived and manufactured based on biological principles of design. The emphasis on utilizing native biopolymers such as collagen and biocomposites in addition to manufacturing procedures operating at the molecular and nanoscale levels has intensified and this approach permits a more accurate control of chemical and physical properties in the final macrostructure [163, 164]. The resultant structures are designed to be similar to the naturally occurring ones. Electrospun materials are examples that offer fibrous constructs with dimensions resembling native extracellular matrices and therefore offer the same chemical and physical properties [165].

Naturally occurring biomatrices such as marine shells and sponge skeletons with wide-ranging structural similarities and chemical homologies to human extracellular matrices and whole tissues have thus far been identified as candidates in our search to find scaffolding materials for bone tissue engineering. Certain species of these marine animals has been applied to the regeneration of human bone and the main source of natural skeletons has been natural and converted corals due to its crystallographic, structural, and chemical similarities to native human bone [158,159,160, 166]. The application of the hydrothermal processing has allowed natural skeletons to be used directly as a scaffold for growing cells into tissues and ultimately in the creation of new bone tissue [166]. Since then, nacre seashells, marine sponges [167], echinoderm spines [168], and the invertebrate marine skeletons of hydrozoans and cuttlefish [169], and others have been examined. Their micro- and macrostructures with suitable pore sizes and its interconnected structural networks are the logic behind the use of these structures as clinical bone-graft materials, which permit easy pathways for sustaining and organizing the growth of bone tissue. Furthermore, diatoms possess natural hierarchical structure making them ideal as functional materials. Despite the lack of comprehensive examination, the regenerative potentials of diatom biogenic silica (such as the transformation of diatom frustule into nanoparticles) have been considered for bone tissue engineering [170].

The development of cost-effective and efficient techniques to synthesize various calcium phosphate phases from biogenic natural materials has received significant amounts of research efforts. Due to their unique compositions which is mainly calcium carbonate as well as their architectures, a number of marine animals such as seashells, nacre, Mediterranean mussels, sea urchins, cuttlefish, and sea corals have been suggested for possible conversions to calcium phosphates bioceramics such as hydroxyapatite and tri-calcium phosphate (α- and β-TCP) for biomedical applications [158,159,160, 166]. Likewise, the skeletal ossicles from sea stars (Pisaster giganteus) have also been studied since they can offer an ideal architecture together with chemical and physical properties conductive to bone restoration [171].

2.3.1.3.1 Coral Skeletons in Tissue Engineering

Used extensively in dentistry, orthopedics, craniofacial surgery, and neurosurgery as a bone replacement due to their combination of open porosity, good mechanical properties, and their proficiency to generate chemical bonds with soft tissues and bone in vivo, natural coral exoskeletons have the best mechanical properties of all the porous calcium-based ceramics in general (Fig. 2.5) [158,159,160]. More importantly, their rates of resorption have been observed to be the same as the formation of new host bone tissues. The excellent biocompatibility and mechanical properties of coral were the result of its organic composition. The composition, abundance, and conformation of the organic matrices are responsible for the successful biological integration of coral with human host [172].

Fig. 2.5
figure 5

Natural coral structure displaying interconnected pores and its architecture [106]

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The exoskeleton of marine madreporic corals is used as a starting material to produce natural coral graft substitutes. The use of coral for dental applications were first attempted by researchers in 1929 and since then corals were recognized and examined in animals during the early 1970s and later in humans in 1979 as a possible candidate as bone graft substitutes. The structure of Porites, a commonly used coral, is comparable to that of human cancellous bone and its initial mechanical properties are also similar. Coral grafts, even though not osteogenic or osteoinductive, are suitable to serve as carriers for growth factors and permit the attachment, growth, differentiation, and spread of cells. Natural coral exoskeletons have been observed to be an excellent bone grafts substitutes if they are utilized in a proper fashion. Consequently, it is also important to select the most appropriate coral so that its resorption rate is identical to the rate of bone formation at the implantation site. In addition, using the composite approach, studies were carried out to combine calcium phosphate converted from corals with polymeric matrix such as polyvinyl acetate and polylactic acid to produce a novel composite bone substitute with improved mechanical properties and functionality [173, 174].

As a consequence of their calcium carbonate structure with high dissolution rates, the application of coral skeletons for tissue engineering and general routine orthopedic surgery has been limited to external fixation devices and not fitting for strictly load-bearing applications. Sol–gel coating techniques on the other hand can be used to improve the strength of corals and as a result enables coral to be applied more frequently at various skeletal locations (Fig. 2.6) [1, 3, 11, 12, 106, 158,159,160].

Fig. 2.6
figure 6

SEM images of the coral structure a before and b after conversion to hydroxyapatite showing surface morphological changes. c Micrograph showing sol–gel coating effectively obliterates the surface meso- and nano-pores, while leaving the macropores intact [180]

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Corals, either in their natural or hybridized synthetic forms, offer immense opportunities in bone tissue engineering. Enhancements in osteogenesis is observed when coral skeleton is combined with in vitro expanded human bone marrow stromal cells at levels greater than those recorded with pure scaffold or scaffold including fresh marrow [175]. Clinical results from orthopedic and maxillofacial surgical cases using in vivo large animal segmental defect revealed there is a complete re-corticalization and formation of a medullary canal with mature lamellar cortical bone and onlay graft for contour augmentation of the face, giving rise to clinical union in a lot of cases [176, 177].

Structural and biomineralization studies of coral can be utilized to enhance the development of new advanced functional materials owing to the unique nanoscale organization of organic mineral and tissue. For instance, the characterization of the ultrastructure of deep-sea Bamboo coral (Anthozoa: Gorgonacea: Isididae) revealed the internodes display bone-like biochemical and mechanical properties. Furthermore, the organic matrix of the coral, which is comprised of an acidic fibrillar protein framework, shows potential as a model for future applications in tissue engineering. The opportunities arising from the use of this Bamboo coral in tissue engineering are still not fully exploited and understood. The growth of both osteoblast and osteoclast are supported by this organic matrix. It has been suggested that blood vessel implants can be fabricated using the collagen matrix (gorgonin) of this coral due to its exceptional bio-elastomeric properties. Quinones can be utilized to cross-link and harden collagenous gorgonin proteins and the resultant product resembles closely to human keratin. Lessons could be provided by the mechanisms used in the synthesis of gorgonin and the way it interacts with the mineralization process for the manufacture of synthetic collagen-like materials [178].

As mentioned previously, utilizing marine-derived calcium carbonate exoskeletons for long-term load-bearing applications is not ideal owing to their fast degradation rate. This property, on the other hand, might prove to be potentially useful in drug delivery application where short-term and fast-acting therapy is required. The conversion of the calcium carbonate exoskeleton of coral to the more stable calcium phosphate structures and its derivatives such as hydroxyapatite, TCP, and their mixtures as biphasic apatites in an effort to overcome such limitations [1, 3, 11, 12, 106, 158,159,160]. To produce calcium phosphates and its derivatives from coral, the abovementioned hydrothermal exchange conversion technique is used in which the carbonate component of the coral is substituted by phosphate at temperatures of around 200–260 °C for a period of 24–48 h (Fig. 2.6) [179, 180]. The adjustment of the molar ratio of calcium to phosphate enables various forms of calcium phosphate to be produced.

Compared to other calcium phosphate compositions, hydroxyapatite or TCP are more suitable for drug delivery applications and for bone grafts under certain conditions. Furthermore, owing primarily to its relatively faster dissolution rate, TCP compared to hydroxyapatite have been extensively investigated and applied as bone grafts [1, 3, 11, 12, 106, 158,159,160]. This controllable dissolution rate also renders TCP more ideal for drug delivery applications [160]. In order to produce TCP from calcium carbonate exoskeletons, a calcium to phosphate molar ratio of 1.5 is required during the hydrothermal conversion process. Most importantly, the amount of time used for the conversion will determine if carbonated TCP is produced. A complete transformation will occur if the conversion took longer than 48 h, while carbonated TCP is produced when the conversion time is less than 24 h [3, 11, 12]. It should also be mentioned that the conversion time is also dependent on the amount of exoskeleton being converted.

A double-stage conversion method was also developed with the intention to improve the mechanical properties and bioactivity of the converted calcium phosphate. The first stage of the process involved the complete conversion of calcium carbonate exoskeletons until 100% calcium phosphate is attained. During the second stage, a calcium phosphate nanocoating derived from the sol–gel approach is deposited directly to cover the meso- and nanopores within the intra-pore material, while maintaining the large pores [1, 3, 11, 12, 160, 180]. The bioactivity is improved due to the nanograin size and hence large surface area that enhances the reactivity of the nanocoating (Fig. 2.6).

2.3.1.3.2 Marine Shells (Nacre)

The assessment of nacre and their capacity to be applied in bone tissue engineering has been conducted using sheep, rabbit, and human models [181, 182]. In human patients, fresh woven bone bonded itself completely and penetrated the nacre implant, accompanied by the increased activities of osteoclasts and osteoblasts. However, the degradation and resorption of nacre are restricted even though their acceptance in vivo is affirmed, and this could hinder their application within calcified tissue that requires speedy self-regeneration [181, 182].

As an improbable source of biomaterial, the use of the outer nacreous layer of a certain species of mollusc shell for clinical applications and more specifically the engineering of new bone has proven to be valuable. Mollusc shells can provide deeper insights into the complexities of biomineralization such as the interactions between proteins and minerals and how it can be governed and regulated.

First discovered in the early 1990s [183] and later in 2000s [181, 182], the scientific basis concerning the ability of nacre to combine with bone tissue was closely scrutinized and the results demonstrated that skeletal cells were activated by the nacre, and this in turn induced bone formation as well as providing structural support in a human clinical trial [183]. The mechanism behind the methods used by nacre to directly induce human cells towards the creation of new bone could be explained using the idea that a “signaling” biomolecule is accountable for the regulation of cell-mediated biomineralization, and this is common to both nacre and bone tissues of vertebrates (Fig. 2.7).

Fig. 2.7
figure 7

SEM image of plate-like structure of nacre (a) and combined in vitro culture of human bone marrow stromal cells (b). Nacre chips of various sizes can clearly be seen. The cell mass is stained red for ALP secretion (a primary marker of bone formation) [106]

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Different to any other biomaterials, nacre from the pearl oyster (Pinctada maxima) is capable of inducing osteogenesis and bone formation from latent osteoprogenitors along an endochondral pathway that consists of a cartilage tissue intermediary phase. Commonly known as the mother of pearl, nacre is the lustrous aragonitic inner layer discovered on molluscan shells in taxa such as abalone and mussels and the powdered form have been used in several studies as composite scaffolds [184, 185]. Nacre contains both inorganic and organic component, is mechanically tough, rapid biodegradable, non-immunogenic, and does not result in detrimental physiological effects. These properties are attributable to the plate-like design and the organic content of the nacre. In addition, the organic shell matrix consisted of polysaccharides, proteins, and glycoproteins that act as templates for the mineralization of calcium carbonate.

The “water soluble matrix fraction” or WSM of nacre has been recognized to induce the formation of new bone directly and demonstrated to increase bone mineral density in an ovariectomized mouse model of osteoporosis [186,187,188]. They have also shown to enhance the secretion of cytoplasmic Bcl-2, a key inhibitor of apoptosis. The molecules from nacre matrix displayed a decrease in bone resorption by reducing osteoclast metabolism [187]. Mobile signal transmitters, involved in the biological control of mineralization and serving as an initiator and inhibitor of calcium carbonate crystallization at the growing front of mineralization, has been suggested to disperse into solution based on the evidence available. They have also been displayed to induce differentiation of surrounding latent osteoprogenitor cells [189]. Recently, a study was attempted to improve the release of WSM and calcium ions from natural nacre through intensive crushing into nanometer particles to fully expose the organic bioactive factors of the nacre crystals [185].

The WSM of nacre can be broken down into several fractions containing amino acids of different compositions and sizes. Size exclusion high performance liquid chromatography (HPLC) of the WSM has revealed protein fractions rich in alanine and glycine, with certain biochemical influences on human fibroblasts that controls the differentiation and proliferation of cells [190]. Peptides are widespread in the nacre matrix and certain fractions have been demonstrated to give rise to specific responses from cultured osteoblast cells. For example, protein fractions with molecular weights less than 1 kDa up-regulated ALP secretion, while secretion decreased for high molecular weight fractions. One hundred and ten molecules in the 100 to 70 Da range consisting of glycine-enriched peptides with structural similarities and high affinities for each other have been identified when comprehensive examinations of bioactive low molecular weight molecules were carried out. An increase in human fibroblast cell ALP expression was explicitly displayed by a highly defined matrix protein known as p10 and in the 10 kDa size range [191]. Furthermore, a soluble p60 protein conglomerate extracted from decalcified nacre possesses adequate bioactivity on MSC and 3T3 cells to induce the secretion of mineral nodules [106, 158]. Novel single proteins such as PFMG3 in addition to p10 and p60 have also been discovered and all sourced from the pearl oyster Pinctada fucata and these proteins were found to improve crystallization of calcium carbonate in vitro [106, 158].

Moreover, numerous proteins have been recognized within the different species of nacre and a number of them have been suggested to play a role in regulating bone tissue. The oyster Crassostrea gigas is an example where four unique proteins were identified from proteomic nacre analysis and believed to aid in shell mineralization, with roles in osteogenesis and with structures comparable to endogenous human proteins. However, there were uncertainties regarding whether nacre proteins were the main cause of osseointegration despite suggestions that nacre was the one responsible. An in vivo study has revealed bone-to-nacre apposition and bonding did occur directly due to the availability of favorable surface chemistry that is rich is phosphorous provided by the nacre which is ideal for the recruitment and attachment of osteoblast and osteoclast and matrix synthesis. However, it failed to stimulate an in vivo osteogenic response [192]. Another study has hypothesized that the rationale behind the excellent bonding between nacre and bone tissue was the presence of an organic matrix which generates a favorable surface charge for optimal biological association once the nacre is implanted. This creates a new interfacial microenvironment that produces many functional associations with the surrounding bone tissue [193].

2.4 Regeneration of Hard Tissues: Bone Infections and Drug Delivery

Wound contamination as well as post-operative infections during surgical intervention or after implantation in maxillofacial surgery and orthopedics are concerns widely recognized by the medical community that could jeopardize the osseointegration process and can result in serious clinical problems. Antibiotics, as a result, are often provided to the patient as a precaution and they are administrated intravenously or in the oral form.

Moreover, bacterial infections associated with the use of implantable prosthetics and devices in dentistry and orthopedics are complications that happen frequently and by far the most common. Regarded as an opportunistic pathogen, pseudomonas aeruginosa causes indwelling device-related infections, particularly in catheters. In cystic fibrosis patients, pseudomonas aeruginosa is the main cause of mortality and morbidity. Staphylococcus aureus infections on the other hand results in serious infectious complications such as severe deep-seated infections (osteomyelitis, endocarditis, and other metastatic infections), septic-thrombosis, and/or severe sepsis. Staphylococcus aureus and Staphylococcus epidermidis strains including methicillin-resistant staphylococcus aureus are some of the biofilm infections that are acquired through basic hospital and surgical intervention. In order to defend against biofilm infections, gaining a deeper understanding into the molecular bases of biofilm formation is vital [160].

Numerous mechanisms and interactions are involved during the process in which bacteria adheres to surfaces of materials such as implants and prosthetics, which includes physicochemical, cellular, and molecular. In addition, other factors can influence the complexity of this phenomenon such as the environment where adhesion takes place and the presence of bactericidal substances or serum proteins. Moreover, the characteristics of the bacterial itself and the surface properties of the materials where adhesion will take place also play an important role throughout the process [194].

Given the fact that different adhesives may have been used by bacteria on various surfaces (i.e., different acceptor), limitations have been places on some of the hypothesized models and theories as physical interaction between the bacteria and surface of material is taken into consideration while disregarding the biological attributes of adhesion. Adhesins are certain bacterial structure responsible for adhesive activities and they regulate the adhesion between different cells as well as cells and abiotic surfaces. From a surface charge, chemistry, and hydrophobicity perspective, the charge of the cell wall and of the substrate are affected by the pH and ionic strength of the medium and this in turn governs their interaction [160].

As discussed above, the development of bone infection is based on the formation of a bacterial biofilm where the bacteria differentiate into sessile forms from planktonic that protect some bacterial cells, which can be released from the biofilm after the termination of antibiotic treatment. Osteomyelitis in orthopedics occurs primarily in tissues surrounding the infected area. The return of bony sequestrum is common and their manifestation is possible given the fact that the existence of damaged bone prevents healing and surgical intervention is needed for their removal.

A widely accepted model of the physiology and pathology associated with osteomyelitis is the devascularization of bone once it becomes infected and bacteria is concealed by the resultant sequestered segments of necrotic cortical bone. The chronicity of osteomyelitis is caused by this sequestered, necrotic, infected bone. The infected sections are encapsulated by pus and granulation tissue and the growth in devascularization of the surrounding bone is driven by bacterial toxins and enzymes through increase in intraosseous pressure. As the infection becomes chronic, relatively avascular fibrous tissue replaces the granulation tissue which envelops the infected area. Motivation to create new reactive bone known as involucrum occurs within the surrounding tissues as well as within the periosteum permeative MSCs that surround the sequestered, necrotic, infected bone [160].

Once the fibrous and bony encapsulation transpires, antibodies and antibiotics must traverse this involucrum and relatively avascular fibrous tissue to reach the microorganisms. The microorganism is also isolated from the defenses of the host as a result of the very effective effort by the human body to quarantine the host from the infection. The infection cannot be eliminated and becomes chronic as soon as this pathological stalemate takes place. This situation is the starting point for surgical intervention and removal of necrotic sequestra. As soon as they are no longer isolated, antibiotic therapy completes the treatment of chronic osteomyelitis by destroying the microorganisms [195]. Thus, models that utilize elimination of infection as a criterion for success along with histological findings of new bone formation, inflammation, sequestration, and intraosseous bacteria, combined with cultures, will provide the most accurate technique for assessing diseases.

Osteomyelitis can be treated with immobilization and antibiotic therapy utilizing a number of medications such as fusidic acid, flucloxacillin, tobermycin, gentamicin, or vancomycin. The main issue related to the application of antibiotics is to make sure its activity and release is maintained for an extended period post operation [196]. Furthermore, high dosage of some antibiotics has been shown to be nephrotoxic and ototoxic as reported by previous studies. The loaded dosages for most controlled release systems are typically high and for that reason, the systemic exposure of antibiotic in urine and blood is a primary safety concern [160].

Thus, a more ideal form of treatment is to deliver antimicrobial agents locally and in a targeted manner. Various approaches have been devised on either controlling and/or preventing bacterial infections. The development and deposition of multi-functional and multi-layered nanocoatings such as nanocomposite coating or nanolaminate on medical implants and devices that have an effect against microbial adhesion or viability will create new opportunities in the prevention of device-related infections. This approach will in turn alter the surfaces of medical devices physically, chemically, or biologically to make the surface free of microbial adhesion.

Several synthetic and natural biomaterials and bioceramics have been suggested as potential candidates to be added to nanocoatings and nanocomposites as biodegradable drug delivery systems [1,2,3]. On the other hand, manipulating these materials into suitable shape with sufficient microporosity and inserting them into bone defects of various sizes was found to be more difficult. The most ideal material for bone repair and slow drug delivery systems is calcium phosphate [1,2,3]. Along with their ease of production and drug carrying capacity, they can supply both calcium and phosphate during dissolution. Using a carrier device such as a composite coating on an implant is one of the most effective method of achieving targeted delivery of antibiotics at a specific site and slowly releasing the appropriate dosage. The search for a more efficient method to deliver antibiotics without the toxicity of systemic antibiotics and complications connected to long-term intravenous access has been ongoing. During the last decade or so, numerous studies have been conducted on both commercial and experimental drug delivery vehicles based on calcium phosphate [1,2,3].

The delivery of antibiotics due to their wide area of applicability has become a major emphasis in the treatment and prevention against infection during surgical interventions and post-operative period. Commercial products based on acrylic polymers or bioceramics such as bioactive glass or calcium sulfate are presently available in pellet form and can be used as slow drug delivery devices. In spite of this, their effectiveness is restricted, and they are either resorbed by the natural physiological process of the body rapidly or being non-resorbable and a second surgery needs to be carried out for their removal due to issues associated to their size and shape, chemical composition, and dissolution rates.

Recently, it has been demonstrated that marine materials such as corals, Foraminifera, and marine shells with specific microspherical design offer desired functions for the delivery of antibiotics [158,159,160]. The synthesis of a composite consisted of a biodegradable polymer and coralline-derived calcium phosphate loaded with clinically active substance such as gentamicin have been proposed to act as a coating on metallic implants and fracture fixation devices for the prevention of implant-associated infections [158,159,160]. The use of biodegradable polymer is advantageous due to their ability to degrade over time. It has also been suggested that they could enhance drug stabilization and increase the drug encapsulation efficiency of the composite coating. In addition, it is possible to engineer the rates of drug release to suit the treatment by regulating the interconnectivity and pore sizes of the coralline-derived calcium phosphate particles.

Investigations have been carried to determine the relationship between calcium phosphate particles embedded into the biodegradable polymeric matrix such as PLA and the release profile of loaded antibiotics such as gentamicin [160]. The release kinetics of gentamicin was found to obey the Power law Krosmeyer Peppas’ model with mainly diffusional process via several different drug transport mechanisms. Furthermore, the release profiles were reported to display an early burst stage followed by a continuous rate of release with significant antimicrobial activity against S. aureus even at high concentration of bacteria. Even after four weeks of drug release, the growth of bacteria is controlled by the composite coating.

As discussed previously, meso- and nanopores are found within calcium phosphate materials converted from coral. Corals converted to calcium phosphate will contain meso- and nanopores as previously discussed, and throughout the initial loading phase, pharmaceutical drugs such as gentamicin will seep into the pores and the porous network. They will also completely coat the surface of the converted calcium phosphate during the process. Consequently, the term “progressive dissolution process” can be used to describe the three-stage drug dissolution process of composite coating comprised of drug-loaded particles embedded within a biodegradable polymeric matrix.

During the first stage, the immediate dissolution of surface-bound drugs in physiological environment caused by concentration gradient and exposed outer surface area results in an initial burst of antibiotics and presumption can be made that the diffusion of drugs from polymeric matrix surface regulates this release [197]. The rate of dissolution during the second stage is governed by the internal diffusion of drugs infused inside the matrix. It is also conceivable diffusion could occur within the porosities created during the synthesis of the matrix. In comparison to the first stage, the second stage is somewhat slower owing to drug conveying through primarily from the large pores and surface areas of the particles. The diffusion continues with the release of antibiotics from narrow pores of the meso- and nanopores of the particles to the matrix and then to the environment. It is important to note that only dissolution of the drugs takes place in the second stage and there is no degradation of the particles. The use of calcium phosphate could in several ways slow down the release of antibiotics through its internal micropores. The final stage is the terminal release phase, and it involves the degradation and dissolution of the particles together with the gradual deterioration of the polymeric matrix. This results in the slow release of the added drugs or the minerals into the environment. Throughout this stage, there is pore growth due to both mass loss by polymer degradation and micropores joining or amalgamating to create larger pores.

2.5 Concluding Remarks

The birth of nanotechnology has generated novel ideals for the manufacture of synthetic bone-like calcium phosphate nanomaterials and nanocomposites. Due to its ability to imitate the structure and composition of the bone mineral hydroxyapatites, synthetic calcium phosphate is the perfect material when deciding a bone replacement. Furthermore, the accessibility of calcium phosphate nanomaterials has undoubtedly opened new prospects for the development of superior biocompatible coatings for implants and high-strength nanocomposites. It has been well documented that despite calcium phosphate share similarities in terms of chemistry and composition with human bone, its mechanical properties are far from being ideal. This issue can be addressed through the creation of a nanocomposite that combines the bioactivity of calcium phosphate with other nano- and microscale materials as a secondary phase, and by utilizing this approach, it is feasible to design and produce nanocomposites with mechanical properties much closer to those of natural bone.

Acquiring greater understandings into bioceramics currently used as bone grafts or tissue engineering scaffolds could significantly contribute to the design of new-generation implantable devices. In addition to synthetic calcium phosphate, the application of bioglasses is of vital importance in the restoration of physiological functions by helping the human body to promote tissue regeneration or to heal when used as body-interactive materials. Their application can be further examined in the advancement of next-generation bioactive glasses that can include specific drugs and biogenic materials in an effort to improve their capabilities and functionality. The field of tissue engineering in recent times has been directed to seize the advantage of the combined application of three-dimensional bioglass or ceramic structures and the use of living cells to engineer neotissue to the damaged site of the patient. Currently, practicable and productive strategies have been targeted at merging the acquired knowledge applied to the field of cell growth and differentiation of osteogenic cells with a relatively traditional approach such as bioglass implants.

During the last decade, the marine environment has been extensively explored from biomimetics to soft and hard tissue engineering and controlled drug delivery. The use of biomimetic approach can create promising outcomes for applications in tissue engineering of skeletal tissues. In the future, advanced biomimetic scaffolds must be able to adapt and undergo the ever-changing needs of developing tissues. It is expected that the production of biomaterial scaffolds will interact with surrounding cell population at the macro (bio-functional level), micro (architectural level), and nano/meso (at the contact interface) levels. Nanofabrication using biological principles of assembly and design is still in its early stages and the application of this bio-inspired nanofabrication for tissue engineering is an exceptional approach which has enormous potential to enhance scaffold design that contours to the physicochemical environments with the capacity to micro-evolve. Presently, there is an obvious need for better tissue engineering scaffolds that possess greater natural bio-responsive environments beneficial to guiding the natural processes of regeneration, which can be extremely intricate and dynamic in space and time.

Without doubt, bacterial infections are complications most frequently associated with the use of implantable medical devices such as fracture plates and dental and orthopedic implants. Numerous strategies have been proposed to lessen the problem related to biofilm infections and they are based on either controlling and/or preventing bacterial infections. The development of multi-functional nanocomposite coatings with surface properties that have an effect against microbial viability or adhesion is one promising approach in the prevention of device-related infections. The search is ongoing to discover a more efficient and less costly way to deliver antibiotics to fight against bacterial infections without the complications associated with long-term intravenous access as well as the toxicity of systemic antibiotics. In the future, research will focus on a process known as gene editing where certain genes of bacteria that are responsible for antibiotic resistance are deactivated and thereby disabling biofilm resistance. This may improve the capability of existing antibiotics to treat infections involving biofilms. Surfaces that would slow-down the formation of biofilm or preventing the colonization of pathogenic bacteria through anti-bacterial or anti-biofilm agents would represent a major improvement in the long-term survival of medical implants and implantable devices.