Surface Transformations of Bio-Glass 45S5 during Scaffold Synthesis - Assignment Example

Surface Transformations of Bio-Glass 45S5 during Scaffold Synthesis Introduction Bioactive glasses possess biocompatible properties and are known to exhibit strong inter- surface bonding with bones. This bioactivity is attributed to the hydroxycarbonated apatite (HCA) layer which forms of their surface. Notably, these layer bear large similarity to a bone’s mineral part. Tissue bonding rate is therefore dependent of the rate of HCA formation; a process that follows reaction sequences between the implanted material and boundary tissues as well as physiologic fluids. Hench et al. () described a 3-step HCA formation mechanism when bioactive glass comes is brought into contact with physiologic fluids. These include; ion exchange, dissolution and precipitation. Ion exchange takes place at the surface of the bioactive glass whereby cations like Na+ and Ca2+ from the bioactive glass interchange with H+ from the surrounding solution. Dissolution on the other hand, results from breakage of Si–O–Si bonds through hydroxyl ions action. The resulting hydrated silica formed on the bioactive glass surface undergoes rearrangement through neighboring silanols polycondensation, and as a result, a silica-rich gel layer is formed. Calcium and phosphate ions released from the bioactive glass alongside those from the solution create a calcium–phosphate-rich layer on the surface through precipitation. In physiological fluid, hydroxycarbonate apatite layer resembling the bone mineral develops on Bioglass 45S5 surface. The layer adsorbs the surrounding tissue’s collagen attracting osteoblasts creating favorable environment for bone regrowth. Consequently, bioglass is considered an osteoinductive material (Shea & Mooney, 2009). Nonetheless, there is a limitation in its brittleness where the glass cannot be solely used in healing of extensive bone defects. This is overcame by using the bioglass to form a composite scaffold with poly (d, l-lactide) (PDLLA); a biodegradable polymer. Bioactive glass has a weight composition as shown; 45% of SiO2, 24.5% of Na2O, 24.5% of CaO, and 4% of P2O5 with sodium ion being the first to leach. The Na+ is exchanged with protons (H+) from the solution around it. Speedy Na+/H+ exchange raises the pH of the solution and Bioactive Glass silica network commence the process of dissolution. As Si-O-Si bridges rupture, there is formation of silanol groups which gradually undergo condensation to form a silica-rich layer. It is this silica-rich surface layer that facilitates migration of Ca2+ and PO4 from Bioactive Glass bulk to the surface leading to formation of calcium phosphate (Ca-P) rich layer. Ca-P layer crystallization and OH- and CO3 integration from the solution leads to CA formation on the Bioactive Glass surface. After this occurs, tissue fibrillar collagen integrates with the layer of HCA, and as the recruitment of osteoblasts occurs, there is an established biological bond between glass and the newly formed bone. This is what makes Bioactive Glass osteoinductive, and hence its usage in dental and orthopedic applications, more particularly, in promotion of regrowth of bones is tiny, non-load bearing structures, for instance, inner ear bones. However, as already mentioned, Bioactive Glass has low fractural toughness; a problem that is overcame through creation of composite Bioactive Glass-polymeric structures, which are usable as scaffolds for filling bones. The resultant composite scaffolds are porous and hence allow nutrients and oxygen to get to the osteoblasts seeded within, and have a prime role in bone tissue engineering (Bouhadir & Mooney, 2012). The resultant structures temporary withstand loading at the injury site. They gradually undergo bio-degradation as surrounding tissues are regenerated. Bg addition into polymeric scaffold incites regrowth of the host tissue, favor cell proliferation as well as components extracellular production and improve bone defect filling in comparison to scaffolds which do not have Bg. Scaffolds and its preparation HAP (Ca10 (PO4)6 (OH)2) is a major mineral part of bones. It can be reabsorbed into bone tissue due to the osteoclasts action. Various researchers have reported that HAP can directly bond with host’s bone (Cook, 2008). Additionally, HAP is an osteo-conductive material and can increase bone regeneration rate. It can further be used as filler in reinforcement of scaffold’s mechanical properties. The present research considers a 3-dimensional, porous HAP scaffold. The scaffold was prepared and characterized using a three-step procedure (Novikok, 2008; Roudier et al., 2010). First step involves gelation of HAP solution to produce a hydrogel with cations (Ca2＋, Sr2＋, and Ba2＋) (Sutherland, 2011). Secondly, is freezing the product from step 1 and then finally drying by lyophilization to result into a 3-dimensional, porous sponge. Figure 1: Scaffold Preparation MATERIALS AND METHODS Materials Sodium alginate of low viscosity obtained as so were Hydroxyapatite (HAP) and calcium chloride. HAP of powder size was availed as follows; 98.1 wt. % smaller than 45 m, 1.4 wt. % measured between 45 and75 m, while 0.5 wt. % were larger than 75 m. also availed was Strontium chloride and calcium gluconate. Dulbecco's Modified Eagle's medium powder was also availed. Fetal bovine serum was also obtained and so was rat osteosarcoma UMR106 cell line; which is an osteoblastic cell line. Preparation conditions Preparation Conditions for Alginate/HAP Composite Scaffolds Cross-linker Type & Freezing Concentration (Moles) Temperature Alginate/HAP ratio CaCl2 (0.030) －10C 100/00, 75/25, 50/50 CaCl2 (0.030) －40C 100/00, 75/25, 50/50 CaCl2 (0.030) －80C 100/00, 75/25, 50/50 CaCl2 (0.030) Liquid N2 100/00, 75/25, 50/50 CaCl2 (0.010) －40C 100/00, 75/25, 50/50 SrCl2 (0.010) －40C 100/00, 75/25, 50/50 Ca-Gluconate (0.010)  40C 100/00, 75/25, 50/50 Preparation of alginate/HAP composite scaffolds In table above is a list of preparatory conditions for pure alginate and alginate/HAP composite sponges. In order to successfully prepare pure alginate sponge (alginate/HAP: 100/00), 1.544 g of alginate powder was thoroughly dissolved and mixed in 40 mL of double-distilled water. A homogenizer at 26,000 rpm was used for a period of 3 min. 10 mL of cross-linker solution was then added to the alginate solution and continuously stirred using the homogenizer at 26,000 rpm for 3 min. Alginate solution concentration was maintained at a constant 3% (w/v) for all prepared scaffolds. Afterwards, the alginate gel solution was left to remain undisturbed for a period of 30 min before being quenched. To minimize heat transfer during the process of quenching, special copper boxes were deigned to provide uniform temperature distribution around this alginate solution. The quenching process started through pouring of the alginate solution into the small copper box. The copper box was then capped and moved to already pre-cooled shelf of a bigger copper box to attain the desired temperature in a freezer. The alginate gel solution was then frozen for 6 hours and lyophilized overnight with the help of a freeze dryer at 0.1 to 0.2 torr. The freeze drying temperature was - 45C. When liquid nitrogen is used as the quenching medium, gel solution within the small, capped copper box was directly dipped into liquid nitrogen bath for a 1 min period. Two kinds of scaffolds were prepared on basis of alginate and HAP weight ratio. The desired HAP powder amount was first dispersed fully in 40 ml double-distilled water using homogenizer at 26,000 rpm for a period of 3 min. afterwards, 1.55 g of alginate powder was truly homogenized using dispersed HAP solution at 26,000 rpm for a period of 3 min. after this, 10 mL of cross-linker solution was introduced into the alginate/HAP solution and fully mixed using the homogenizer. The alginate/HAP gel solution was consequently molded, frozen, and then dried using the procedure described above. Bio-active glasses Since bone consists of large amounts of hydroxyapatite (HA), Ca10 (PO4)6(OH) 2, HA and related calcium phosphates (CaP) (e.g., β-tricalcium phosphate) have been considered to develop scaffold materials for bone regeneration (Shea & Mooney, 2009). Due to their close chemical and crystal resemblance to the mineral phase of bone HA and CaP exhibit excellent biocompatibility. The close similarity of hydroxyapatite to the mineral component of bone, which is stable in the body, results however in the lack of biodegradation of HA in the body, which is generally an undesirable feature for tissue engineering scaffold materials (Gepstein, Weiss, & Hallel, 2010). For example, a recent clinical report on a 6-7 year follow-up study has confirmed that implanted crystalline HA is not biodegradable, remaining in the body for extended periods with no visible signs of biomaterial reabsorption. Bioactive silicate glasses (e.g. 45S5 Bio-glassR) with compositions in the system SiO2-Na2O-CaO-P2O5, having 950 oC) during scaffold fabrication and that the mechanically strong crystalline phase can transform to a biodegradable, amorphous calcium phosphate at body temperature and in a biological environment. This transformation enables the two normally irreconcilable properties, i.e. mechanical competence and biodegradability, to be combined in a single scaffold. This discovery promises to go some way towards the scaffold optimization and its clinical application. It was found with the 75/25 alginate/HAP composite scaffold. HAP was distributed evenly on the pore wall surfaces of the sponges. The same morphology was also found with the 50/50 alginate/HAP composite scaffold. This suggests that the addition of HAP to the polymer solution did not change the morphology of the scaffold. In the quenching process, the bottom surface of the scaffold was exposed near the interface between the alginate solution and the bottom of the freezer, where the cooling rate is the fastest. In this area, many ice-crystal nuclei were formed. At such a rapid rate of cooling, however, the heat of crystallization was extracted and the nuclei did not have enough time to grow, thus preventing the formation of large ice crystals. As a result, the bottom surface of the scaffolds had the smallest round, interconnected pores. The boundaries of the pore walls were blurred because the pore structure was not well developed due to the fast cooling rate. The top surface of scaffold, on the other hand, was exposed near the interface between the alginate solution and the air, where the cooling rate is the slowest. In this area, ice-crystal nuclei had enough time to grow; therefore, they produced large pores and a well-interconnected pore structure (Figure 3B, 3E, and 3H). The cooling rate at the midsection of the scaffolds, however, was different from those at the top and bottom surfaces. The different cooling rates at the top and bottom surfaces induced a temperature gradient in the midsection of the scaffolds. In this section, the pores were elongated and indicated the direction of crystal growth. It was also found that the pore size of scaffolds decreased as the quenching temperature decreased (Table 2). Table 2. The Effect of Freezing Temperature on the Pore Size of Scaffolds at the Bottom, Top, and Midsection 48 Alginate/HAPa content (weight/volume) 100/0 75/25 50/50 Temperature Bottom Top Midsection Bottom Top Midsection Bottom Top Midsection 10C 187  5 191  8 212  15 195  5 193  7 196  22 71  2 159  12 210  12 40C 122  5 160  6 155  7 129  5 162  10 158  9 119  3 126  4 163  11 80C 79  3 126  5 143  6 42  2 98  6 109  5 36  1 133  9 178  10 Liquid N2 < 10 < 10  < 10 < 10  < 10 < 10  a3% alginate, cross-linker: 0.03 M CaCl2. Conclusion It was concluded that cooling rate is faster at lower quenching temperatures, smaller ice crystals are produced in the alginate/HAP solution. The pore sizes of the pure alginate and 75/25 alginate/HAP scaffolds decreased as the freezing temperature decreased (Table 2). The differences in pore size were statistically significant for the bottom pores of the pure alginate scaffolds and for the top, bottom, and midsection pores of the 75/25 alginate/HAP scaffolds (p < 0.05). The variation in pore size of the 50/50 alginate/HAP scaffolds may have been caused by adherence of the HAP powder to the bottom during mixing or by some sedimentation during gelling. The average of all pore sizes for all the prepared scaffolds was approximately 150 m, which is close to the result from other investigators using thermally induced solid-phase separation to fabricate the porous materials. A smaller pore size was obtained by using a quicker freezing process. The pore size of the scaffolds was smaller than 10 m on the top and bottom surfaces when the sponges were quenched by directly dipping into liquid nitrogen. The effects of type and concentration of cross-linking agents on the pore size of the prepared scaffolds are listed in Table 3. In general, the pore sizes of all the scaffolds prepared using 0.03 M CaCl2 solution as a cross-linking agent were greater than those prepared using 0.01 M CaCl2 solution. The differences were statistically significant (p < 0.05) for the tops and bottoms of the 100/0 and 75/25 scaffolds. In general, the pore sizes of the scaffolds prepared from 0.01 M of CaCl2, SrCl2, and Ca-gluconate solutions were not significantly different at the tops or midsections. In regeneration of bones, scaffold faces the biggest challenge in its mechanical properties when it comes to replacement of large bone defects. In such cases, load transmission is a challenge. Despite recent technological advances, multiple issues still need to be addressed in clinical application of scaffolds. The question of scaffolds mechanical reliability, vascularization induction and tailored degradability are of great concern. Incorporation of biomolecules like growth factors the objective of accelerating local bone healing shows immense promise although it still requires extensive research. References Shea, L. D, & Mooney, D. J. (2009). Engineered bone development from a pre-osteoblast cell line on three-dimensional scaffolds. Tissue Eng., 6 (4), pp. 605-617. Gepstein, R, Weiss, R. E, & Hallel, T. (2010). Bridging large defects in bone by demineralized bone matrix in the form of a powder. A radiographic, histological, and radioisotope-uptake study in rats. Journal of Bone Joint Surg. Am, 69 (3), pp. 984-992. Sutherland, I. W. (2011). Biomaterials: novel materials from biological sources. New York: Stockton Press. Cook, W. (2008). Alginate dental impression materials: chemistry, structure, and properties. J Biomed Mater Res, 20 (12), pp. 1-24. Klock, G. et al. (2010). Biocompatibility of mannuronic acid-rich alginates. Biomaterials, 18 (2), pp. 707-713. Novikok, L. (2008). A novel biodegradable implant for neuronal rescue and regeneration after spinal cord injury. Biomaterials, 23 (13), pp. 3369-3376. Roudier, M. et al. (2010). The reabsorption of bone-implanted corals varies with porosity but also with the host reaction. J Biomed Mater Res, 29 (7), pp. 909-915. Bouhadir, K. H. & Mooney, D. J. (2012). Synthesis of hydrogels: alginate hydrogels. New York: Academic Press; 2002:653-662. Read More