Bioactive Molecule-loaded Drug Delivery Systems to Optimize Bone

Current Protein & Peptide Science

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REVIEW ARTICLE ISSN: 1389-2037 eISSN: 1875-5550

Bioactive Molecule-loaded Drug Delivery Systems to Optimize Bone Tissue Repair

Impact Factor: 2.441

BENTHAM SCIENCE

João Augusto Oshiro Jr.1, Mariana Rillo Sato1, Cassio Rocha Scardueli2 , Guilherme José Pimentel Lopes de Oliveira2, Marina Paiva Abuçafy1 and Marlus Chorilli1,* 1

Faculdade de Ciências Farmacêuticas, UNESP-Univ Estadual Paulista, Araraquara-Jaú, Km 1, Araraquara, 14800903 SP, Brazil; 2Faculdade de Odontologia, Universidade Estadual Paulista (UNESP), Humaitá, 1680, Araraquara, 14801-385 SP, Brazil

ARTICLE HISTORY Received: February 13, 2017 Revised: March 17, 2017 Accepted: March 21, 2017 DOI: 10.2174/1389203718666170328111605

Abstract: Bioactive molecules such as peptides and proteins can optimize the repair of bone tissue; however, the results are often unpredictable when administered alone, owing to their short biological half-life and instability. Thus, the development of bioactive molecule-loaded drug delivery systems (DDS) to repair bone tissue has been the subject of intense research. DDS can optimize the repair of bone tissue owing to their physicochemical properties, which improve cellular interactions and enable the incorporation and prolonged release of bioactive molecules. These characteristics are fundamental to favor bone tissue homeostasis, since the biological activity of these factors depends on how accessible they are to the cell. Considering the importance of these DDS, this review aims to present relevant information on DDS when loaded with osteogenic growth peptide and bone morphogenetic protein. These are bioactive molecules that are capable of modulating the differentiation and proliferation of mesenchymal cells in bone tissue cells. Moreover, we will present different approaches using these peptide and protein-loaded DDS, such as synthetic membranes and scaffolds for bone regeneration, synthetic grafts, bone cements, liposomes, and micelles, which aim at improving the therapeutic effectiveness, and we will compare their advantages with commercial systems.

Keywords: Bioactive molecules, drug delivery systems, osteogenic growth peptide, bone morphogenetic protein, tissue repair. 1. INTRODUCTION Bone healing is a complex process and difficult stage of reconstruction in medicine [1]. Although this tissue has the potential to self repair, this process can be committed in major defects that are associated with the absence of bone walls [2]. Critical defects are bone losses that do not heal spontaneously, due to the inability of bone progenitors to migrate long distances and difficulty in obtaining a complete local revascularization, impairing tissue maintenance through lack of nutrients and oxygen. Therefore, successful bone healing is intimately connected with vascularization, stability at the defect site, and with the presence of bone cells. However, critical defects require biomaterials with properties that help the bone formation and can be expressed by different mechanisms of action (osteoinduction, osteocondution, and osteogenesis) [3]. Several studies show different strategies to help bone formation, with diverse biomaterials and techniques, and consequently, miscellaneous results [4]. The search for *Address correspondence to this author at the Faculdade de Ciências Farmacêuticas, UNESP-Univ Estadual Paulista, Araraquara-Jaú, Km 1, Araraquara, 14800-903 SP, Brazil; Tel: +551633016998; Fax: +551633016900; E-mail: [email protected] 1875-5550/17 $58.00+.00

treatments with greater predictability of bone formation in critical defects is the target of new studies [4, 5]. Three essential elements of bone regeneration (osteogenesis, osteoinduction, and osteoconduction) are necessary for successful bone healing, acting separately or in synergy. The presence of all these biological elements involved in bone formation permits higher success rates [2]. The osteogenic characteristic of progenitor bone cells (osteoblasts and osteocytes) shows a higher activity for bone formation, with viable bone cells directly linked to production or maintenance of bone matrix [6]. The autogenous bone is the single graft that presets this property of bone formation at the moment [7]. Similarly, some molecules exhibit an activity that stimulates the differentiation of mesenchymal or undifferentiated cells into osteogenic cells, which are capable of producing bone tissue (osteoinduction) [3]. These molecules can be found in autogenous bone, or isolated and incorporated in different bone substitutes [8]. Furthermore, many grafts or biomaterials possess only the characteristics for maintenance of skeleton for bone formation, allowing the presence of blood (osteoconduction) and consequently cells and nutrients. Osteoconduction examples are inorganic bone and scaffolds (without inductive molecules) [2].

© 2017 Bentham Science Publishers

Bioactive Molecule-loaded Drug Delivery Systems

Current Protein and Peptide Science, 2017, Vol. 18, No. 8

In bone grafting in medical or dentistry areas, the autogenous bone is considered the gold standard [9], as it shows all properties related to bone formation. However, due to bone removal from a donor area associated with greater morbidity, limited tissue availability, additional surgical pain, and advance of biotechnology, some biomaterials can be of interest, especially when associated with bioactive molecules, leading to similar properties and results to the autogenous bone [10-12]. The use of bioactive molecules for bone formation has arisen and gained attention; when associated with their binding proteins, they regulate cell functions (proliferation, migration, and differentiation) of bone cell lineages. Thus, several molecular mediators, such as proteins and peptides, can optimize the bone formation process [13, 14]; among them, bone morphogenetic proteins (BMPs) and osteogenic growth peptide (OGP) stand out. 1.1. Bone Morphogenetic Protein Proteins extracted from bone tissue can be used to induce the formation of bone and cartilage. In 1965, Urist identified an osteoinductive protein by demonstrating that mineralized bone matrix implanted in muscular tissues induces the formation of cartilage and bone tissue [15]. These findings confirmed the presence of bioactive factors in demineralized bone, as referred to BMPs [16]. However, it was only in 1988 that Wang et al. reported the isolation of BMP from bovine bone tissue [17]. Consequently, its DNA was cloned, and by using peptide or amino acid sequence homology to human database, it was considered a new member of the transforming growth factor β (TGF-β) family [18, 19]. BMPs are released during repair and remodeling processes, due to degradation of inorganic bone by osteoclasts. Therefore, in vivo studies with animals showed promising results on the role of BMP in bone synthesis without association with osteoconductive biomaterials [20]. So far, 20 different types of BMPs with distinct characteristics and roles have been described; nevertheless, they all share osteoinductive characteristics that induce differentiation of mesenchymal

B

cells into osteoblasts [21]. However, only BMP-2 and BMP7 have been approved for clinical use [15]. Although BMPs are multifunctional proteins, recombinant human BMP-2 (rhBMP-2) plays a major role in the repair of fractures, and the absence of this protein causes development abnormalities in bone and impairs the healing of fractures [22]. BMP-7 stimulates the synthesis of erythropoietin, a hormone that controls the production of erythrocytes from precursors in the bone marrow. However, so far, it is still not totally understood how BMPs regulate the formation and repair of tissues [15]. It is known that, being soluble molecules and acting as local signaling proteins, BMPs are involved in inducing specific markers of osteoblast cell differentiation [23] and can induce a differentiated phenotype in some types of cells [24]. Therefore, BMPs can regulate growth, differentiation, chemotaxis, and apoptosis, and play pivotal roles in the morphogenesis of a variety of tissues and organs. Nonetheless, the activity of BMPs can be controlled at different levels, intracellularly through Smad proteins (specific receptors at the surface of the cell that transduce signals to transcription factors that activate specific genes) [25], and tissue-specific transcription factor (basic helix-loop-helix factor and its binding sequence (E-box)). Extracellularly, many protein inhibitors can bind BMPs and inhibit their binding to cell surface receptors [25, 26]. All these mechanisms act in a controlled and limited manner, and BMP expression is only observed in localized site (Fig. 1). In vivo studies in animals showed that BMPs were able to induce bone and cartilage, when implanted in ectopic tissues [18], or improve bone formation [27, 28]. Although BMP potential to induce bone formation has been demonstrated in preclinical studies, the findings in clinical trials are limited but show positive results for bone formation [29-33]. However, although positive, many factors are uncertain and limit the clinical use of BMPs, such as fast degradation, high costs, high doses, osteolysis, high inflammation, and ectopic bone formation. To limit fast liberation and consequently lesser action, BMP's can be associated to release vehicle that

BMP

MP

BMPreceptors

BMPreceptors P

P

P P

Smad

Phosphate grouping

P

Smad

Osteoblast

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Phosphorylation Smad P

Smad

DNA

Mesenchymal undifferentiated cell Fig. (1). Possible mechanism of action of BMPs in bone formation.

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difficult the degradation and can be released slowly, increasing its biological activity [34, 35]. Currently, the use of high doses is to compensate BMP short half-life [36]. Thus, some side effects may be present, such as bone reabsorption (osteolysis), edema, and ectopic bone formation that need extra medical treatments [37, 38]. The beneficial effects of BMPs in regeneration or reparation of bone defects are known; however, the mechanisms and their potential collateral effects remain undiscovered, thus cautious use and more studies are needed. 1.2. Osteogenic Growth Peptide OGP is a sequence of 14 well-preserved amino acids (Ala-Leu-Lys-Arg-Gln-Gly-Arg-Thr-Leu-Tyr-Gly-Phe-GlyGly). Its sequence is similar to C-terminal histone H4 (89102), and was isolated from the blood during bone remodeling or bone marrow regeneration [14]. The pentapeptide in its cleaved form (OGP 10-14) has demonstrated an even greater potential for bone synthesis. An analysis of the structure indicates that its activity is related to the phenolic group of tyrosine (Tyr10) and the aromatic ring of phenylalanine (Phe12) [39]. OGP is present physiologically in mammalian serum in micromolar concentrations; its small size and linearity make it susceptible to proteolysis, and the low concentrations of OGP in vivo suggest the presence of a circulating protein binding partner for OGP. This specific or non-specific protein provides protection against proteolysis and acts as a “system” that releases low concentrations of OGP into the circulation [40-44]. It is believed that the protein binding partner of OGP at higher concentration is α2-microglobulin (α2M). After cleavage of complex α2M-OGP, the peptide is cleaved proteolytically [41, 42]. Similar to BMPs, OGP has been shown to be a potent osteoinducer due the stimulation of the differentiation of undifferentiated mesenchymal cells into osteoblasts, preventing the formation of condroblasts and adipocytes [40]. This osteoinductive effect was demonstrated in studies with mesenchymal stem cells derived from bone marrow with OGP, as well as OGP (10-14) [41-44]. The mechanisms by which OGP induces osteoblast differentiation and increases osteoblastic activity are unknown. Different studies demonstrate several mechanisms of action of OGP on bone formation (Fig. 2). It has been suggested that the stimulation of osteoblastic differentiation may occur via the RhoA/ROCK pathway [45] and CDK2/cyclin A [46]. The effects of OGP on osteoblastic activity may be due to the upregulation of MAPK and ERK 1/2 pathways [47]. OGP regulates the expression of various TGF-β such as TGF-β1, TGF-β2, and TGF-β3 [48]. In vitro studies have shown that OGP regulates cell proliferation and alkaline phosphatase (ALP) activity in osteoblastic cell lines through self-regulatory feedback mechanism [49, 50]. Another possible mechanism of action of OGP is the induction of osteoprotegerin (OPG) production by osteoblasts [51]. This action interferes with the RANK/RANKL/OPG pathway, involved in the formation of osteoclasts. Preclinical studies have shown that OGP improved the healing of bone fractures [52] and the osteodistraction of

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long bones [53]. Furthermore, it was also demonstrated that OGP promoted the maintenance of bone tissue structure in ovariectomized rats, indicating a possible use of this peptide in osteoporosis therapy [54-57]. 2. APPLICATIONS OF BIOACTIVE MOLECULELOADED DRUG DELIVERY SYSTEMS TO OPTIMIZE BONE TISSUE REPAIR Despite the promising results of isolated administrations of bioactive molecules, they are often unpredictable because their biological half-life is short, and they show low longterm stability, are high cost treatment, and have low tissue specificity, requiring administration of high doses, which in some cases may exceed the potential carcinogenicity dose [58]. These disadvantages led researchers to seek new strategies for the administration of these substances. One solution to these problems is the development of vectorization systems capable of transporting the active substance, improving its distribution in specific tissues, and limiting the side effects by preserving the efficacy of the active compound [51, 55]. Recently, in medical and other fields, the interest in new technologies for drug delivery has increased. Thus, the combination of bioactive molecules with DDS presents further advantages compared with conventional devices, such as decreased toxicity, prolonged half-life in the bloodstream, gradual and controlled release of the drug, safe (no local inflammatory response), and adequate (small dose requirement) administration, targeting specific sites, and the ability to incorporate lipophilic and hydrophilic substances [58-60]. The functionalities of these DDS (protection against degradation, targeting, controlled release, easier cell penetration) are due to their internal structure and surface properties that give them better physicochemical properties, such as shape, composition, molecular weight, identity, purity, stability, and solubility [59], allow an increased therapeutic efficacy of many drugs by modifying the pharmacokinetics or biodistribution of the active molecules, have lower costs, and maintain the concentration of these factors within the therapeutic range over a longer period of time [60]. Thus, the osteoprogenitor cells migrate to the target site and differentiate into osteoblasts, resulting in a major impact on current therapies for bone tissue regeneration [20, 58]. These physicochemical properties allow the development of DDS to be used in different approaches in clinical practice, such as synthetic membranes and scaffolds, synthetic grafts, bone cements, liposomes, and micelles. Therefore, in this section, we mainly introduce different approaches of DDS with OGP and BMP aiming at therapeutic effectiveness in bone formation. 2.1. Synthetic Membranes and Scaffolds The regeneration of a tissue needs to be oriented to result in a new and functional physiology. The guided bone regeneration (GBR) technique assists the restoration of bone tissue by preventing competition between bone tissue cells and soft tissue cells [61].



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Osteogenic Peptide Osteogenic Growth Growth Peptide

Regulates Regulatesexpression expression TGF-b TGF-b

Mesenchymal MesenchymalCell Cell

Active Active ERK1/2 MAPK and ERK1/2 MAPKand

Regulation Regulation transcription transcription factors factorsof of osteoprogenitory osteoprogenitory cells cells

Alkaline Alkaline Phosphatase Phosphatase

Osteoprotegerin Osteoprotegerin (OPG) (OPG)

PO 3434-- OO OPOP OO

RANKL RANKL

OO

OPG OPG Mineralization Mineralizationof of the bone bone matrix matrix

Osteoprogenitorcell cell Osteoprogenitor

Increasednumber numberofof Increased Osteoblasts Osteoblasts

RANK RANK

Osteoprogenitor Osteoprogenitor cell cell

Osteoclasts Osteoclasts

Bone Formation Formation Bone Fig. (2). Possible mechanisms of action of OGP in bone formation.

The principle was first described in 1959 by Hurley and contributors in an experimental treatment that aimed for a mechanical barrier, to isolate the tissues in a spinal fusion surgery [62]. However, this pioneering study did not immediately lead to a broad clinical application in patients. The clinical potential of GBR was only recognized in the early 1980s when Karring et al. evaluated synthetic membranes for periodontal regeneration in several clinical trials [63]. A few years later, based on the promising results of these studies, Dahlin et al. tested a polytetrafluoroethylene membrane to isolate the connective tissue of defects created in bone tissue of rat mandibles, and defects without membrane were used as control group. The results showed that the control group presented lower bone growth (2.2 mm) when compared to the defects treated with the membrane (3.8 mm). Another factor observed in the control group was the presence of soft tissue cells within the bone defects. These results reveal the importance of the use of synthetic membranes in the regeneration of bone tissue [64]. Commercially, synthetic membranes can be made of polytetrafluoroethylene, which is classified as nonresorbable, or polylactic acid, collagen, or polyglatin, which are classified as resorbable [65]. The non-resorbable membranes require a second surgical procedure to remove them, which may compromise the success of the procedure, since the second surgical intervention may disturb newly formed tissues [66]. In addition, it may

result in patient discomfort, increased cost, and post-surgical infection [67]. Resorbable membranes, however, have the advantage of eliminating this second surgical phase, as well as the trauma to the neoformed tissues. Nevertheless, they present invasion of soft tissue cells resulting from membrane degradation and instability at the surgical site (non-fixation) [65]. To overcome the disadvantages of resorbable and nonresorbable membranes, researchers are looking to developing membranes from new materials [68], with characteristics such as physical barrier, cell affinity (osteoconduction), and promoter of bone growth (osteoinduction) (Fig. 3) [51, 6971]. For this purpose, it is necessary that the membranes are able to incorporate and release controlled bioactive molecules that are capable of stimulating migration and differentiation of mesenchymal cells, such as BMPs, fibroblast growth factor, TGFs, and bone growth peptide among others [71, 72]. Another advantage of these materials is the ability to maintain the therapeutic concentration of these bioactive substances, since BMPs have a low plasma half-life (2-4 hours) when administered alone, they must however be on site for a period of 4 weeks to induce a successful bone regeneration process [72]. Bioactive electrospun fibers based on poly(lactide-coglycolic acid) (PLGA) by immobilizing bone-forming peptide 1 (BFP1) derived from the immature region of BMP-7 were developed by Lee et al. in order to achieve an ideal


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Biomaterial

Bioactive molecules Bone defect

Treatment with Biomaterial

(a)

(b)

(c)

Bone Formation

Fig. (3). Synthetic membrane able to incorporate and release bioactive molecules acting as: physical barrier and as promoter of bone growth (osteoinduction) [69].

barrier synthetic membrane [73]. The membrane was tested in a 4 mm critical-sized calvarial bone defect mouse model. The authors observed that after 8 weeks, the bone volume formed in the defects with membrane was statistically higher than that of the control group. The semi-quantification of bone volume revealed that the area was approximately 20% in the defect-only group and 57.59 ± 15.24% in the group implanted with a membrane. The results suggest that membranes are interesting substances for the treatment of criticalsized bone defects [73]. Pigossi and co-workers [55] prepared bacterial cellulosehydroxyapatite (BC-HA) membranes associated with OGP and OGP (10-14) at the concentration of 10-9 mol L-1 to treat critical-size calvarial bone defects (4-mm diameter) in mice. It was concluded that a high percentage of bone formation was observed after 60 days; however, this effect was associated with BC-HA, and the addition of OGP had low effect [55]. Policastro et al. [42] studied amino acid-based poly(ester urea)s (PEU), degradable polymers functionalized by tethering OGP to tyrosine-based monomer subunits to improve repair of bone defects in rats. The authors observed that after 4 weeks, PEU-OGP resulted in matrix mineralization without inflammatory response. Moreover, migration of blood vessel within the material was observed, demonstrating the osteogenic and angiogenic capacity of these materials [42]. Lopes et al. [74] studied the effect of laser light in bone defects (diameter = 5 mm) filled with BMP on a commercial membrane. The results showed that the association of these techniques improved significantly improved bone formation when compared to the defects without laser therapy [74].

Among membranes for bone regeneration, scaffolds have emerged as the new generation membranes with features suitable for tissue regeneration. These biomaterials have the ability to act as a template that allows the cells to replace it by newly formed tissues. One of the strategies used in the development of these materials is the immobilization of bioactive molecules with osteoinductive capacity in order to get better results [75, 76]. Therefore, to establish more effective treatments of large bone defects, Kolambkar et al. [76] reported the use of an alginate-based hybrid system with rhBMP-2. These authors analyzed the release of bioactive BMP-2 (500 ng) from this system. The results demonstrated that 71.2 ± 3.8 ng was released within 21 days; however, 98.6% of BMP-2 was released with 7 days. Critical-sized (bilateral 8 mm segmental defects) femoral segmental defect in a rat model was used to verified bone formation. They observed that after 12 weeks, the hybrid system with bioactive rhBMP-2 (37.65 ± 2.22 mm3) had significantly more bone volume relative to the system alone (3.96 ±1.40 mm3). These results indicate that these systems can be used for bone repair in a GBR procedure [76]. Lee et al. [77] bound BMP-2 with heparin-binding peptide amphiphile (HBPA) nanofibers, which were introduced in an absorbable collagen material, aiming for bone regeneration in a rat critical-size femoral defect model (5 mm). It was verified that a decrease in the release of BMP-2 occurred when nanofibers contained heparan sulfate (HS) complex. The amount released after 8 days reached approximately 84.0 ± 18.9% of rhBMP-2-HBPA without HS complex, while the release of protein reached 34.5 ± 8.1% with


HBPA-HS. The in vivo bone formation was evaluated after 6 weeks and showed a significant increase in the bone volume in defects treated with HBPA-HS-BMP-2-collagen (20.3 ± 4.2 mm3) relative to the control group of untreated animals (2.4 ± 0.1 mm3), those treated with collagen scaffolds with BMP-2 (8.2 ± 1.7 mm3) or without (3.5 ± 0.8 mm3) it [77]. Aiming to improve the chemical conjugation methods between polymer and bioactive molecules (requiring multistep and complicated procedures), Ko et al. [78] conducted a study in which BMP-2 was efficiently immobilized onto PLGA scaffolds by a single step of polydopamine-mediated coating. The system was transplanted into a mouse model of calvarial bone defect (diameter: 4 mm). The results revealed that after 8 weeks, the group treated with PLGApolydopamine-BMP-2 exhibited enhanced bone regeneration (~ 60%) while the bone formation was ~ 38% with no treatment [78]. Other studies tested OGP as an integral part of tissue engineering and showed conflicting results. The use of OGP and OGP (10-14) in coating of biocellulose membranes did not promote greater formation of bone tissue relative to the control membranes, despite the induction of greater expression of bone formation biomarkers [55]. OGP associated with crosslinked PEU polymer-based scaffolds showed biocompatibility and enhanced mechanical properties and bioactivity (reabsorption and revascularization) when these scaffolds were inserted in the dorsum of rats [56]. It also showed that the addition of OGP in PLGA scaffolds improved the healing of segmental bone defects in long bones [57]. Further studies are needed to identify the best parameters of using OGP as a coating of membranes and associated with biomaterials. The concentration of this peptide, the ideal type of scaffold or membrane, and the type and concentration of cells that should be used with OGP as a component of tissue engineering remained to be determined. After obtention of these data, clinical trials that evaluate the effect of OGP in bone repair can be performed. However, the three-dimensional (3D) printing technology represents an alternative to fabrication of GBR membranes, which traditionally used solvent casting and it is difficult to make a freeform membrane with a given thickness, pore size, and external shape, factors that contribute to the amount of bone formed, since they influence the release of the active substance and the adhesion/penetration of cells into the material [79]. Thus, Shim et al. [80] investigated a 3D printingbased PCL/PLGA/β-TCPGBR membrane that incorporated BMP-2 (5 µg mL−1). This system can be used to control drug release. In vitro studies demonstrated that BMP-2 release reached ~ 25.5% in 24 h and was sustained for 28 days. To evaluate bone formation, the membrane was tested in an 8mm critical-sized calvarial bone defect rabbit model. The results showed that after 8 weeks, the bone volume formed in the defects with membrane loaded with BMP-2 represented~ 60%, which was statistically higher than that of the control group (15.24%) [80]. These results demonstrated that better conditions in the manufacturing process (3D printer use) can help the development of biomaterials with optimized specifics (thickness, external shape and number of pores standardized), sustained release of bioactive molecules and high capacity to bone formation.


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2.2. Synthetic Bone Grafts One of the most significant advances in biomaterials over the last few years has been in the field of bone synthetic graft substitutes. They have been extensively studied in order to be substituted to autografts, due to some disadvantages such as poor bone quality, inadequate amount of bone available, and a possible immunogenicity that limits the use of these grafts. Thus, synthetic bone graft has become a potential material for clinical applications in different medical areas. There are characteristics that synthetic bone grafts should have, such as being of various shapes and sizes with suitable mechanical properties to be used in different sites, biocompatible, osteconductive, and preferably being resorbable and replaced by new bone formation [81] (Fig. 4). Synthetic bone grafts should act as scaffolds for bone cells, providing bone growth inside the material. Therefore, the scaffold must have adequate mechanical properties and be highly porous, and these pores have to be interconnected. Several physicochemical features are considered fundamental for successful tissue engineering such as surface chemistry and roughness, mechanical properties, topography, and interfacial free energy [82, 83]. Reabsorption of biomaterial is linked to several factors, such as particle size, porosity, and crystallinity [84]. Particles with nanometric size are reabsorbed faster than micrometric size ones, and highly crystalline structures are more resistant to resorption than amorphous ones. Bone growth proteins, such as OGP and BMP, marked the history because they were two of the first materials to be produced by recombinant genetic technology. These materials have given rise to a new generation of synthetic and bioactive bone graft substitutes. Shuqiang et al. [57] evaluated the effect of locally applied OGP incorporated into a synthetic bone graft, PLGA scaffolds, in healing bone defects in rabbits (1.5 cm segment defect was made in the right radius using a small saw). OGP was incorporated into PLGA, and was released during 7 days. The in vitro result revealed that 90% of peptide total quantity was released after a week. After 12 weeks, the rate of development of bone bridging was significantly higher in the group with OGP than in the control group. The experiments concluded that OGP release from the synthetic bone graft at the site of bone defect resulted in faster and more complete bone healing, and suggested that the porous structure and pore size of material were appropriate for carrying OGP [57]. Song et al. [85] studied a vancomycin-bearing synthetic bone graft that delivers rhBMP-2 in rats. They developed a poly(2-hydroxyapatite methacrylate)-nanocrystalline hydroxyapatite (pHEMA-nHA) synthetic bone graft with 3 µg rhBMP-2 preabsorbed per graft. The experiments were done with 5-mm femoral defect in SASCO-SD male rats (289-300 g of body weight). They observed that after 12 weeks, pHEMA-nHA-vancomycin without rhBMP-2 resulted in partial bridging of defect, and full bridging with mineralized restoration was achieved with rhBMP-2. The complete composite repaired the 5-nm rat femoral segmental defects. The success of this strategy requires that pHEMA-nHA effectively retains and releases vancomycin and rhBMP-2 without


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NEW BONE GRAFTS TO BONE REGENERATION

ADVANTAGES Synthetic Bone Grafts + Bioactive Molecules (OGP or BMP)

. Various shapes and sizes with suitable mechanical to be used in different sites; . Biocompatible; . Osteconductive, and preferably being resorbable and replaced by new bone formation

Shorter treatment time and lower cost = GREATER ADHERENCE OF THE PATIENT TO THE TREATMENT

Fig. (4). Advantages of the bone grafts with bioactive molecules to bone regeneration.

compromising the structural integrity of the graft or its ability to promote bone cure [85]. Smith et al. [86] analyzed an animal evaluation of synthetic bone graft, a paste of chitosan glutamate and hydroxyapatite. Cranial defects of 8-mm diameter were made in rat calvaria, and the defect evaluation was studied with pure paste and paste with 40 µg of BMP-2. Rats were sacrificed at intervals during 18 weeks. Calvaria containing a defect were harvested, and bone mineral density was determined by energy X-ray absorptiometry. This study demonstrated that a paste of chitosan glutamate and hydroxyapatite could be used to deliver BMP-2, and the bone mineral density and histology data confirmed that this paste containing BMP-2 formed mineralized bone and showed osteoblastic activity [86]. Tagil et al. [87] investigated a synthetic bone graft composed of a tricalcium phosphate hydroxyapatite scaffold that can be used with BMP-7 (25 µg). They performed a 6-mm femoral nailing, and the defects were grafted after 4 week. After 11 weeks, the bones were explanted for evaluation with radiography, micro-CT, and infrared spectroscopy. Isolated scaffolds without BMP-7 did not heal any defects, whereas treatment with BMP-7 had greater volume of highly mineralized bone and higher volume of fraction. It was concluded that synthetic scaffold with BMP-7 can heal a critical size defect [87]. 2.3. Bone Cement Cement is a compound made of two materials, a liquid and a powder form, which when mixed make a moldable paste that can be adapted to variable surfaces of imperfect bone structures [88, 89].

Cement has several attractive advantages compared to other materials, such as the ability to promote an adjustable fit in the fixation of the prosthesis, easy handling, adherence to the hard tissue, hypoallergenic or non-carcinogenic characteristics and increased resistance to loading by uniformly distributing the tensions between the prosthesis and the bone [90-92]. Moreover, it is able to induce osseointegration and act as a release system of several molecules, which have a synergistic effect that increases the effectiveness of the material in bone repair (Fig. 5) [93-95]. One of the most promising molecules that can be carried by cement is the osteogenic peptide, such as OGP and BMP, which are used as an interesting alternative because they have a broad spectrum of activity and binding to specific sites of extracellular matrix proteins and can be transported with other drugs [96]. Studies have evaluated the in vivo behavior of OGP and BMP in different routes of administration, such as systemic administration [96] and topical administration [57]. This analysis showed an important result for the route of administration of these molecules, revealing a more effective therapeutic effect when applied locally, which highlights the importance of materials such as bone cement, which can be used as carrier of the biomolecule, allowing its retention in the defect site for prolonged periods of time with better results than via the intravenous route [42]. Kroese-Deutman et al. [97] studied the osteoinductive capability of bone calcium phosphate (Ca-P) cement loaded with rhBMP-2. The cement discs were loaded with 30% of rhBMP-2 in vitro and implanted subcutaneously in rabbits for 2 and 10 weeks. Thereafter, the histological analysis revealed bone formation in rhBMP-2-loaded cement discs with



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Bone Cement (Liquid + Powder + Bioactive Molecuels)

Cure time

Attractive advantages Moldable paste

. Promote an adjustable fit in the fixation of the prosthesis. . Increase of the resistance to loading prosthesis by uniform distributing the tensions between the prosthesis and the bone. . Act as a system of release of molecules, having a synergistic effect to increase the effective of the material in bone repair.

Fig. (5). Advantages of the bone cement with bioactive molecules to bone formation.

pores filled at around 18% after 10 weeks. The fraction of rhBMP-2 retained in the discs in subcutaneous sites was less than 10% after 10 days, probably too low to induce bone formation. This suggests that although rhBMP-2 is a powerful osteoinductive material in other locations, it is less suitable for subcutaneous administration. The pattern of fluorescent labeling showed that bone formation started at the cement surface and then proceeded into the center of pores. The scaffold was stable for 10 weeks, which indicates that the material is a suitable carrier material for bone formation [97]. The in vivo release kinetics of rhBMP-2-loaded PLGA/Ca-P cement was studied by Ruhe et al. [92]. rhBMP2 was radio labeled and encapsulated or adsorbed onto the surface of PLGA microparticles. PLGA microparticles were prepared at high and low molecular weight (HMW PLGA and LMW PLGA, respectively). The adsorption efficiency was 93% for LMW and HMW microparticles, and the fraction of rhBMP-2-loaded microparticles in cement was 54%. Release of rhBMP-2-loaded composites was assessed by scintigraphic imaging during 4 weeks, and the in vivo release kinetics showed release without initial burst and displayed a linear release from 3 to 28 days. The retention of rhBMP-2 was 16% higher for each formulation and time period than that measured by conventional ex vivo counting. The authors affirmed that rhBMP-2 release from the composites was slowed down by an interaction of rhBMP-2 with cement after its release from PLGA microparticles. Therefore, this system can be considered as a sustained slow release vehicle, and the retention of rhBMP-2 can be changed according to need [92]. Another group of researchers [98] worked with rhBMP2-loaded porous cement that was pretreated with albumin

and studied the release in vitro and in vivo. They thought that the pretreatment of the ceramic with albumin prior to rhBMP-2 loading would result in weaker binding between the peptide and the cement and enhanced rhBMP-2 release. These researchers chose albumin for pretreatment because it is the most abundant body protein and its limited specific activity is known. Therefore, the pretreatment of albumin would result in release of the growth factor without changing the properties of the material. Each side of the discs was carefully loaded with 12.5 µL of rhBMP-2. Albuminpretreated cement discs showed a 30% release after 24 h and 22% without albumin. The release kinetics in the rat ectopic model showed 20-30% retention after 4 weeks. Although albumin pretreatment of the cement resulted in an increased initial release of rhBMP-2, the in vivo release kinetics was similar with non-pretreated cement. 2.4. Liposomes Liposomes are formed by spherical structures consisting of one or more phospholipid bilayers (natural or synthetic polymers) around the inner aqueous compartment, which diameter can range from 20 nm up to 1 µm [99-103]. Liposomes are widely used as drug carriers because they promote controlled drug delivery, improve their pharmacokinetic properties and bioavailability, minimize the adverse effects observed in conventional therapies, allow the incorporation of hydrophilic drugs into the inner compartment and hydrophobic ones in the lipid membranes, are biodegradable, biocompatible, and non-toxic, and can be produced on a large scale [104]. Furthermore, the similarity between plasma membrane and phospholipid bilayer structures of liposomes allow a better interaction of the particles with cells and tissues of the organism [105]. Fig. (6) shows


A

Oshiro Jr. et al.

Charge and Polymer Stabilised Liposome

PEGylated lipids

Hydrophilic Drugs +

-

-

+ Charged lipids

-

Hydrophobic Drugs

Stimuli Responsive Agent

+ Hydrophilic Drugs

Imaging Agent Conjugated Drug

Carbohydrate Antibody

B

Ta r g e t e d Liposome

C Peptide

Protein

Small Molecules

Theranostic Liposome

Fig. (6). Schematic representation demonstrated by Lombardo et al. [106] of three types of liposomal drug delivery systems: Charge and polymer stabilized (A), targeted (B), and theranostic (C) liposomes [106].

the schematic representation demonstrated by Lombardo et al. [106] of three types of liposomal drug delivery systems. With the commercialization of Doxil® [107], the first nanostructured drug approved in 1995 by the US Food and Drug Administration (FDA), feasible strategies have been developed with liposomes in the cosmetic and pharmaceutical industries, mainly for the delivery of cytotoxic drugs, genes, and vaccines [108], as well as bone regeneration [109]. In bone regeneration, bone growth peptide and proteins OGP and BMP have been the targets of investigation for therapeutic intervention, as they have shown promising results in vivo and in clinical studies. Thus, the aim of these peptide and proteins loaded liposomes was to develop effective bone regeneration [110]. Park et al. [111] compared liposome-mediated and adenoviral gene transfer for the generation of autologous BMP2-producing bone marrow stromal cells that aim to treat critical size defect of rat femur (diameter of 6 mm). In the 6 weeks after transfer of the BMP-2 gene, the liposome group showed a complete healing of the critical size bone defects, whereas the adenoviral-mediated transfer of BMP-2 cDNA showed bone healing within 4 weeks, showing that both are suitable methods for the healing of critical size bone defects in rats. None of the control groups revealed bone healing, even after 8 weeks. However, considering the ease of preparation, no limitation DNA size, lower immunological and safety problems, liposomes have proven to be more advantageous than any other vector [111]. Matsuo et al. [112] prepared magnetic liposomes with incorporated rhBMP-2 and

evaluated bone formation in a bone defect model in rats (5 mm segment). The results revealed that this system was more effective for treatment of segmental bone defects in rats when compared with control groups, as a unique topical application of magnetic rhBMP-2 liposomes (3.0 µg) incorporated under magnetic induction immediately after surgery showed effective new bone formation. In addition, radiographic assessment over 9 weeks showed significantly higher values (4.8 ± 0.4) and higher bone formation area (12.93 ± 1.80 mm2) than in the control group. This is possibly due to the magnetic induction that enhanced retention of the drug at the implantation site [112]. Ono et al. [113] combined porous hydroxyapatite (HAP) with BMP-2 cDNA plasmid in cationic liposomes as a vector to treat 12-mm diameter bone defects located on rabbit cranium. Bone formation on the cranial defects of four groups of rabbits was histologically evaluated at 3, 6, and 9 weeks after the surgery. The group that received the BMP-2 plasmid with liposomes without HAP showed a marked osseous formation 3 weeks after the operation and a complete bone formation in the cranial defect at 9 weeks, whereas the BMP2 plasmid plus HAP group presented on bone formation after 9 weeks [113]. In a pilot study, Kroczek et al. [114] analyzed the efficiency of four morphogenetic and mitotic osteoinductive proteins, BMP-2, BMP-7, TGF-β, IGF-1, and a liposomemediated gene transfer system on bone regeneration by continuous osteodistraction using Goettingen minipigs. In the histological and radiological analyzes, a complete bone formation was observed after osteoinduction with BMP-2 and


Table 1.


859

Data on bone formation in different critical bone defects using different bioactive molecule-loaded DDS.

DDS

Bioactive Molecule

“In Vivo” Bone Formation

Comments

References

Bioactive electrospun fibers based on poly(lactide-co-glycolic acid) (PLGA)

BMP-7

57.59 ± 15.24% of bone formation after 8 weeks

The membrane was tested in 4mm critical-sized defect.

[73]

Alginate-based hybrid system

BMP-2

37.65 ± 2.22 mm3 of bone formation after 12 weeks

The membrane was tested in bilateral 8-mm segmental defects.

[76]

Heparin-binding peptide amphiphile (HBPA) nanofibers

BMP-2

The membrane was tested in criticalsized defect model (5mm) and showed 20.3 ± 4.2 mm3 of bone formation after 6 weeks.

The amount released after 8 days reached 34.5 ± 8.1%.

[77]

PLGA scaffolds

BMP-2

High bone regeneration (~ 60%) after 8 weeks.

The DDS was transplantated into a calvarial bone defect (4 mm)

[78]

Printing-based PCL/PLGA/βTCP guided bone regeneration (GBR) membrane

BMP-2

8-mm defects presented ~ 60% bone formation after 8 weeks.

~ 25.5% BMP-2 release in vitro at 24 hr and a sustained rate for 28 days

[80]

OGP

A 1.5-cm segment defect was made, and the results showed a rate of development of bone bridging significantly higher in the group with OGP than in the control group after 12 weeks.

90% of peptide released in vitro after 1 week.

[57]

Bone calcium phosphate (Ca-P) cement

BMP-2

Histological analysis revealed bone formation in rhBMP-2-loaded cement discs with pores filled at 18% after 10 weeks of implantation.

Stability of the material during 10 weeks, which indicates that this is a suitable material. More studies are necessary.

[97]

Liposomes

BMP-2

6 weeks after transfer of the BMP-2 gene, the liposome group showed a complete healing of critical size bone defects.

Liposomes were tested in critical size defect of rat femur model (diameter of 6 mm).

[109]

Magnetic liposomes

BMP-2

Statistical significance of radiographic values after 9 weeks (4.8 ± 0.4) and higher bone formation area (12.93 ± 1.80 mm2) were observed.

The liposome was tested in a bone defect model in rats (5-mm segment).

[110]

INFUSE® Bone Graft, Medtronic Spinal and Biologics, Memphis, TN, USA (commercially available in the concentration of 1.5 mg/cc)

BMP-2

12 weeks after implantation, 22.8 ± 10.1% of the defect treated with INFUSE®was filled with new bone, and Ca-P ceramics showed better results with new bone covering 33.9 ± 6.8% of the defects.

INFUSE® Bone Graft used as positive control in in vivo study of bone formation in a bilateral iliac wing defect with a criticalsize diameter of 17 mm in sheep.

[121]

PLGA

BMP-7, with a mean bone formation of 50.1% and 61.0%, respectively. Application of TGF-β, IGF-1, and the liposomal vector had restricted effects on bone regeneration, with quantitative values of 26%, 36% and 30%, respectively. This suggests that the association of osteodistraction with osteoinduction could reduce consolidation and treatment times. 2.5. Micelles Structurally, micelles are defined as thermodynamically stable aggregates of globular amphiphilic molecules or am-

phiphilic block copolymers, with a size between 10 and 100 nm [115, 116]. For applications in the clinical and pharmaceutical fields, polymeric micelles or reverse micelles are particularly relevant DDS [117, 118]. Thus, due to their nanometer size and core-shell structure, micelles have the following advantages: prolonged drug release, incorporation of hydrophilic and hydrophobic drugs, increased bioavailability and solubility, low toxicity, targeting of the specific site of action in a passive or active mechanism, reduced dosage and frequency of


Oshiro Jr. et al.

administration, and lower side effects when compared with conventional drugs, as well as large scale production [119]. Despite these advantages, there are few reports of the use of micelles with bioactive molecules for bone regeneration.

resulting in a major impact on current therapies for bone tissue regeneration.

Ratanavaraporn et al. [120] investigated local suppression of pro-inflammatory cytokines, such as interleukin IL-6 and IL-10, and the effects in BMP-2-induced bone regeneration. Gelatin hydrogels were prepared by incorporating 2.5, 5.0, or 10 mg of immunosuppressive triptolide-loaded micelles and 5 µg of BMP-2 for in vitro and in vivo release studies. The results of in vitro release of triptolide-loaded micelles and BMP-2 from the hydrogels demonstrated a saturated pattern within 7 days. The in vivo studies showed that triptolide and BMP-2 were released simultaneously from the hydrogels over 7 days. When implanted into a criticalsized bone defect in rats (6-mm length), the hydrogels incorporating mixed 2.5 or 5.0 mg of triptolide-loaded micelles and BMP-2 showed a significantly lower number of neutrophils, lymphocytes, macrophages, dendritic cells, or mast cells infiltrated into the defect, and lower expression level of cytokines than those incorporating BMP-2 without triptolideloaded micelles (IL-6: ~900-fold and TNF-α genes: ~80fold). The authors concluded that a lower number of inflammatory cells in defects implanted with the hydrogels incorporating mixed 2.5 or 5.0 mg of triptolide-loaded micelles and BMP-2 enhanced bone regeneration, due to the adequate local modulation of inflammatory responses [120].

João Augusto Oshiro Junior was involved in conceiving ideas and overall management, data collection and evaluation, publication and writing. Mariana Rillo Sato, Cassio Rocha Scardueli, Guilherme José Pimentel Lopes de Oliveira, Marina Paiva Abuçafy was involved in data collection and evaluation and publication writing. Marlus Chorilli was involved in conceiving ideas and overall management.

3. FINAL CONSIDERATIONS

[1]

This review of current research present an account of the innovation that these bioactive molecule-loaded DDS represent in bone repair. Table 1 presents some data of bone formation rate in different critical bone defects, and these results depended on the composition of the bioactive moleculeloaded DDS. These results highlight that DDS may in the future be an innovative approach to mainly replace the use of autogenous graft, which is considered the “gold standard” but has many disadvantages (see introduction), and have the potential to overcome the bone formation capacity of the current treatment with osteoinductive autograft replacement based on rhBMP-2 on a collage sponge (INFUSE® Bone Graft, Medtronic Spinal and Biologics, Memphis, TN, USA) approved by the FDA (see results obtained by Yuan et al. [121] in Table 1), as DDS are able to increase the half-life of bioactive molecules, release them in specific sites, and can be used by consumer/patients resulting in greater amount of bone formed and in less time with a lower cost, benefiting both the surgeon and patient. Therefore, DDS represent a new alternative to conventional treatment of bone defects. The versatility of these materials can be used for the controlled release of bioactive molecules, such as OGP and BMP, offers many advantages such as protection against degradation, targeting, controlled release, easier cell penetration, better therapeutic efficacy by modifying the pharmacokinetics or biodistribution of the active molecules, lower cost and maintaining the concentration of these factors within the therapeutic range over a longer period of time, allowing the osteoprogenitor cells to migrate to the target site and differentiate into osteoblasts,

AUTHORS’ CONTRIBUTIONS

CONFLICT OF INTEREST The authors confirm that this article content has no conflict of interest. ACKNOWLEDGEMENTS This study was supported by Coordination for the Improvement of Higher Education Personnel (CAPES), by São Paulo Research Foundation (FAPESP, Brazil), National Counsel of Technological and Scientific Development (CNPq) and the Programa de Apoio ao Desenvolvimento Científico (PADC) of the School of Pharmaceutical Sciences/UNESP, Brazil. REFERENCES

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