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Sintering effects on mechanical properties of biologically derived dentine hydroxyapatite
October 2002
Materials Letters 56 (2002) 142 – 147 www.elsevier.com/locate/matlet
Sintering effects on mechanical properties of biologically derived dentine hydroxyapatite Gu¨ltekin Go¨ller a,*, Faik Nu¨zhet Oktar b a
Metallurgical and Materials Engineering Department, Faculty of Chemical and Metallurgical Engineering, Istanbul Technical University, 80626 Maslak, Istanbul, Turkey b New Technologies, R&D Center, Marmara University, Ziverbey, Istanbul, Turkey Received 13 February 2001; received in revised form 20 November 2001; accepted 28 November 2001
Abstract In this study, the change in the mechanical properties of hydroxyapatite (HA), which is the main mineral content of bone and teeth, commonly used as an implant material, is tested for different sintering temperatures. Although HA shows great biocompatibility with the human body, its applications are limited to non-load-bearing areas and coatings due to its low mechanical properties. These mechanical properties can be improved substantially by sintering. In this study, naturally produced HA from human teeth is sintered, and density, microhardness measurements and compression tests are performed in order to find the optimum sintering temperature. The results are compared with the synthetically derived HA properties from the literature. D 2002 Elsevier Science B.V. All rights reserved. PACS: 81.05.t; 81.05.Je Keywords: Hydroxyapatite; Sintering
1. Introduction Hydroxyapatite (HA) is the major mineral component of bone and teeth with a chemical formula of Ca10(PO4)6OH2 [1,2]. This material can be hydrothermally converted from coral or synthetically manufactured. HA is well accepted by the human body and has been found to possess osteoconductive properties. The physical properties of HA make it favorable for osseintegration with host bone [1]. Unfortunately, mechanical properties of pure HA ceramics are poor, especially in wet environments. Therefore, HA ceramics cannot be used as heavy-loaded implants, * Corresponding author. Tel.: +90-212-285-6891; fax: +90-212285-3427.
such as artificial bones or teeth. Their medical applications are limited to small unloaded implants such as tooth root substitutes, filling of periodontal pockets, cystic cavities, regions adjacent to implants, spinal fusions, contour and malformation defects and nonunions of long bones. Coating powders with an average particle size of 45 – 125 Am for plasma spray process and low-loaded porous implants are used as a typical graft material with an approximately 50% porosity [2,3]. Synthetic HA is known for its brittleness, revealed by the material’s low fracture toughness value (Klc = 1.1– 1.2 MN m 3/2) [4]. One of the ways to improve the reliability of HA ceramics is to increase its fracture toughness. It can be accomplished by precise control of the microstructure and the use of various reinforcements. In recent years, many rein-
0167-577X/02/$ - see front matter D 2002 Elsevier Science B.V. All rights reserved. PII: S 0 1 6 7 - 5 7 7 X ( 0 2 ) 0 0 4 3 0 - 5
G. Go¨ller, F.N. Oktar / Materials Letters 56 (2002) 142–147
forcements, including particles, platelets, whiskers, long fibres, partially stabilized zirconia and nanoparticles (nanocomposites) have been used in HA materials. The highest values of fracture toughness (Klc = 6.0 – 7.4 MPa m1/2, rf = 175 – 224 MPa) have been achieved for HA containing 20– 30% FeCr alloy long metal fibres [5]. The question remains whether the metal-reinforced HA composites are as biocompatible as pure HA. No results concerning this issue have been presented. In the case of other HA-based composites, Klc was in the range of 2.0 – 3.9 MPa m1/2, depending upon the reinforcement used. An undoubted advantage of the composite approach is an increase in toughness and strength of HA ceramics [2,3]. Tough and strong ceramics are produced with sintering process. ‘‘Sintering’’ is the term used to describe the consolidation of the product during firing. Consolidation implies that within the product, particles have joined together into strong aggregate. The term sintering is often interpreted to imply that shrinkage and densification have occurred. Sintered HA implants are observed to develop cracks [6]. Thus, an optimization of the strength and microstructure of HA implants by a suitable choice of sintering parameters are needed. It has been reported that the critical sintering temperature for calcium phosphates with
143
Ca/P ratio between 1.5 and 1.7 is 1300 jC [7]. Sintering of HA is complicated by the fact that HA is a hydrated phase that decomposes to anhydrous calcium phosphates such as TCP at f 1200– 1450 jC. Decomposition results from dehydroxylation beyond a critical point. For temperatures below the critical point (1300 jC), the HA crystal structure is retained despite dehydroxylation, and the HA rehydrates on cooling. If the critical point is exceeded, complete and irreversible dehydroxylation occurs, resulting in collapse of the HA structure and decomposition. Significant reversible dehydroxylation generally occurs above f 800 jC. After the critical point, a-TCP and h-TCP are often formed. In particular, the molecular volume increase that occurs in the hTCP ! a-TCP transformation seems to be the most deleterious phenomenon for mechanical properties [8].
2. Materials and methods 2.1. Production of HA The HA material used in this study was derived from freshly extracted human teeth. The teeth were irrigated with tap water and soaked in a 1% concen-
Fig. 1. X-ray diffraction diagram of the dentine sample.
144
G. Go¨ller, F.N. Oktar / Materials Letters 56 (2002) 142–147
tration of antiseptic solution to prevent bad odor and contamination of various infectious diseases [9]. Subsequently, the teeth were reirrigated and deproteinized in an alkali solution (1% concentration of sodium hypochloride). Thereafter, samples were reirrigated with tap water again. All samples were calcined at 5 jC min 1 to 850 jC and kept for 5 – 6 h. It was observed that at high temperatures, the dentine material and enamel material separated easily and that nearly 60% of the material was dentine and 40% enamel. Enamel and dentine particles were ground with a blade grinder for 30 s. Samples were poured into sieves placed from coarsest to finest in series. Dentine HA particles between 106 and 150 Am were used [10]. X-ray diffraction analysis of dentine particles has been carried out by Rigaku Rint-100 X-ray Diffractometer. Diffraction results indicated that the dentine material has 100% crystalline structure (Fig. 1). Synthetic HA powder ‘for bioceramics’ with a particle size in the range of 2 – 20 Am (MERCK 2196; Merck, Darmstadt, Germany) was also used for comparison. 2.2. Preparation of samples A die had been designed to prepare samples with a diameter of 12 mm according to a British Standard for
compression tests (no. 7253). The HA powder was weighed to total 3.25 g portions. The prepared portions were pressed at 350 MPa between hardened steel dies (die height 5 mm). In this study, any lubricant like polyvinylalcohol was used. All material were kept in normal atmospheric conditions and kept slightly humid for of a couple days (since dried HA powders would not compact and hold in one piece in the absence of humidity). Pressed samples were sintered in an open atmospheric oven at temperatures 1000, 1100, 1200, and 1300 jC (with the heating rate of + 4 jC min 1) for 4 h. The density of sintered samples was determined by the Archimedes method. Vickers hardness values were obtained using a Shimadzu microhardness testing machine with 200 g load. A universal test machine (Instron 1186) was used for compression tests.
3. Results and discussion The densification of HA as a function of sintering temperature follows a sigmoidal correlation with the attainment of plateau densification levels at 1100– 1300 jC. The plateau temperature, and associated limiting densification level, depend predominantly on the surface area of the HA powder and to a lesser extent on the heating rate and Ca/P ratio [11]. Accord-
Fig. 2. Sinterability of HA at various temperatures.
G. Go¨ller, F.N. Oktar / Materials Letters 56 (2002) 142–147
145
Fig. 3. Variation in compressive strength with sintering temperature.
ing to Suchanek et al. [3], full density can be achieved after sintering at 1400 jC for 3 h. The density of samples for different sintering temperatures obtained in this research can be seen in Fig. 2. The density of HA samples without sintering is qavr = 2.889 g/cm3. It is seen that sintering increases the density of the HA samples. A decrease in density occurring between 1000 and 1100 jC could be attributed to the initiation of h-TCP to a-TCP at that temperature range.
Although a slight increase is observed at 1200 and 1300 jC, the density of samples at 1000 jC is never exceeded. This is probably due to the relatively higher content of h-TCP phase at this temperature. The compression strength and the Vickers hardness values for sintered samples are shown in Figs. 3 and 4, respectively. There is an important increase of compression strength between 1100 and 1200 jC whereas compression strength values is about the same level at
Fig. 4. Variation in Vickers hardness with sintering temperature.
146
G. Go¨ller, F.N. Oktar / Materials Letters 56 (2002) 142–147
1200 – 1300 jC, as seen in Fig. 3. Therefore, it is meaningful to assume that the temperature between 1200 and 1300 jC is the optimum sintering temperature for maximum strength for biologically derived HA. In this study, it is observed that slightly lower results in strength and higher values in density and hardness are obtained compared to studies carried out with the HA powders derived from bone (Table 1). The mechanical test results suggest that the decomposition of biologically derived HA does not start at this temperature range. It is more stable than the synthetic powders commonly used in literature. Three recent investigations have correlated mechanical properties with sintering temperature for HA sintered parts at this range of temperatures. It is stated in literature that the microhardness plateau is similar to the sintering sigmoidal curve, while fracture toughness peaked at 1200 jC and dropped by f 20% at 1250 jC [12]. It is also stated that microhardness and fracture toughness both peaked at the sintering-plateau level, showing a significant decrease about 20– 40% at 1250 – 1300 jC [13]. Moreover, Go¨kbayrak [6] also used biologically derived HA with precisely prepared samples for the BS standard. The maximum strength value he had obtained is given in Table 1. As seen, the compression strength values obtained in this research are close, but density and hardness values are higher than that of Go¨kbayrak data. Wang and Chaki [7] had achieved the maximum density as 2.46 g cm 3 at 1100 jC for synthetic HA. They achieved the maximum hardness also at 1100 jC. It was seen that production of dentine HA samples were very efficient. There was no problem during material preparation whereas samples originated from synthetic HA (2– 20 Am) produced defects [6]. The highest hardness values are obtained at 1300 jC sintering temperature where density and compression strength values are also improved. This result suggests that powders used in this study show better
consolidation properties at higher temperatures. It would be better to perform a chemical analysis to obtain the decomposition rates of the materials at these temperatures for further studies. Ruys et al. [11] found that the apparent density decreased significantly above 1300 jC. The %HA – temperature curve indicated that the critical temperature above which decomposition occurred was 1350 jC. Therefore, 1350 jC is a reasonable value to assign for a nominal value of the decomposition temperature. Additionally, Ruys et al. stated that microstructural development is not dependent simply on sintering kinetics. Furthermore, dehydroxylation effects play an important role. A combination of the sintering and dehydroxylation effects points at two regions of importance in terms of the temperature – strength correlation; first, the attainment of the closed porosity densification level of f 95% at f 1150– 1200 jC and, second, decomposition of HA above 1350 jC. Optimum sintering temperature for synthetic HA is 1100– 1150 jC; above this temperature, it is possible to see phase transformations [14]. The overall results of sintering HA for 4 h at temperature range 1000 –1300 jC are shown in Table 1. Table 1 presents that sintering at 1300 jC makes possible to reach maximum hardness, compressive strength and density for biologically derived HA powders.
4. Conclusions Results indicated that it is possible to obtain dense HA with high strength, hardness and high density by sintering at 1300 jC for 4 h economically, which is suitable to be used as a graft material or tooth substitute. Modifications in sintering conditions and addition of reinforcing materials should be tested by the further studies.
Table 1 Results of sintering at various temperatures Temperature (jC)
qavr (g cm 3), Go¨kbayrak
ravr (MPa), Go¨kbayrak
Hardnessavr (HV), Go¨kbayrak
qavr (g cm 3), Gu¨ltekin
ravr (MPa), Gu¨ltekin
Hardnessavr (HV), Gu¨ltekin
1000 1100 1200 1300
2.46 2.34 2.59 2.48
48.17 22.16 75.20 65.01
85.37 79.20 148.50 130.68
2.992 2.945 2.958 2.961
9.83 14.48 56.30 56.77
2.58 163.1 229.8 751.9
G. Go¨ller, F.N. Oktar / Materials Letters 56 (2002) 142–147
Acknowledgements The authors would like to thank Prof. Dr. Selim Kusefoglu for his contribution during the sintering experiments, and Hakan Ulutas from Devotrans for compression tests.
References [1] J.A. Koempel, B.S. Patt, K. O’Grady, J. Wozney, D.M. Toriumi, J. Biomed. Mater. Res. 41 (1998) 359. [2] W. Suchanek, M. Yashima, M. Kakihana, M. Yoshimura, J. Am. Ceram. Soc. 80 (1997) 2805. [3] W. Suchanek, M. Yashima, M. Kakihana, M. Yoshimura, Biomaterials 17 (1996) 1715. [4] M. Knepper, S. Moricca, B.K. Milthorpe, Biomaterials 18 (1997) 1523. [5] G.D. With, A.J. Corbijn, J. Mater. Sci. 24 (1989) 3411. [6] H. Go¨kbayrak, Production of hydroxyapatite ceramics, MS thesis from the Mechanical Engineering Department of Bog˘azicßi University, 1993.
147
[7] P.E. Wang, T.K. Chaki, J. Mater. Sci.: Mater. Med. 4 (1993) 150. [8] A. Tampieri, G. Celotti, F. Szontagh, E. Landi, J. Mater. Sci.: Mater. Med. 8 (1997) 29. [9] S.V. Joshi, M.P. Srivastava, A. Pal, S. Pal, J. Mater. Sci.: Mater. Med. 4 (1993) 251. [10] F.N. Oktar, K. Kesenci˙, E. Pi˙ßs ki˙n, Int. J. Artif. Cells Blood Substitutes Immobilization Biotechnol. 27 (1999) 367. [11] A.J. Ruys, M. Wei, C.C. Sorrel, M.R. Dickson, A. Brandwood, B.K. Milthorpe, Biomaterials 16 (1995) 409. [12] J.D. Santos, J.C. Knowles, F.J. Monteiro, G.W. Hastings, Development of a glass reinforced hydroxyapatite with enhanced mechanical properties: physical characterisation and in vitro studies, in: T. Yamamuro, T. Kokubo, T. Nakamura (Eds.), Bioceramics, vol. 5, Kobunshi Kankokai, Kyoto, Japan, 1992, pp. 35 – 41. [13] S. Best, W. Bonfield, C. Doyle, Optimisation of toughness in dense hydroxyapatite ceramics, in: G. Heimke (Ed.), Bioceramics, vol. 2, Proceedings of the 2nd International Symposium on Ceramics in Medicine, DKG, Cologne, 1990, pp. 57 – 64. [14] T. Kitsugi, T. Yamamuro, T. Nakamura, S. Kotani, T. Kokubo, H. Takeuchi, Biomaterials 14 (1993) 216.
Materials Letters 56 (2002) 142 – 147 www.elsevier.com/locate/matlet
Sintering effects on mechanical properties of biologically derived dentine hydroxyapatite Gu¨ltekin Go¨ller a,*, Faik Nu¨zhet Oktar b a
Metallurgical and Materials Engineering Department, Faculty of Chemical and Metallurgical Engineering, Istanbul Technical University, 80626 Maslak, Istanbul, Turkey b New Technologies, R&D Center, Marmara University, Ziverbey, Istanbul, Turkey Received 13 February 2001; received in revised form 20 November 2001; accepted 28 November 2001
Abstract In this study, the change in the mechanical properties of hydroxyapatite (HA), which is the main mineral content of bone and teeth, commonly used as an implant material, is tested for different sintering temperatures. Although HA shows great biocompatibility with the human body, its applications are limited to non-load-bearing areas and coatings due to its low mechanical properties. These mechanical properties can be improved substantially by sintering. In this study, naturally produced HA from human teeth is sintered, and density, microhardness measurements and compression tests are performed in order to find the optimum sintering temperature. The results are compared with the synthetically derived HA properties from the literature. D 2002 Elsevier Science B.V. All rights reserved. PACS: 81.05.t; 81.05.Je Keywords: Hydroxyapatite; Sintering
1. Introduction Hydroxyapatite (HA) is the major mineral component of bone and teeth with a chemical formula of Ca10(PO4)6OH2 [1,2]. This material can be hydrothermally converted from coral or synthetically manufactured. HA is well accepted by the human body and has been found to possess osteoconductive properties. The physical properties of HA make it favorable for osseintegration with host bone [1]. Unfortunately, mechanical properties of pure HA ceramics are poor, especially in wet environments. Therefore, HA ceramics cannot be used as heavy-loaded implants, * Corresponding author. Tel.: +90-212-285-6891; fax: +90-212285-3427.
such as artificial bones or teeth. Their medical applications are limited to small unloaded implants such as tooth root substitutes, filling of periodontal pockets, cystic cavities, regions adjacent to implants, spinal fusions, contour and malformation defects and nonunions of long bones. Coating powders with an average particle size of 45 – 125 Am for plasma spray process and low-loaded porous implants are used as a typical graft material with an approximately 50% porosity [2,3]. Synthetic HA is known for its brittleness, revealed by the material’s low fracture toughness value (Klc = 1.1– 1.2 MN m 3/2) [4]. One of the ways to improve the reliability of HA ceramics is to increase its fracture toughness. It can be accomplished by precise control of the microstructure and the use of various reinforcements. In recent years, many rein-
0167-577X/02/$ - see front matter D 2002 Elsevier Science B.V. All rights reserved. PII: S 0 1 6 7 - 5 7 7 X ( 0 2 ) 0 0 4 3 0 - 5
G. Go¨ller, F.N. Oktar / Materials Letters 56 (2002) 142–147
forcements, including particles, platelets, whiskers, long fibres, partially stabilized zirconia and nanoparticles (nanocomposites) have been used in HA materials. The highest values of fracture toughness (Klc = 6.0 – 7.4 MPa m1/2, rf = 175 – 224 MPa) have been achieved for HA containing 20– 30% FeCr alloy long metal fibres [5]. The question remains whether the metal-reinforced HA composites are as biocompatible as pure HA. No results concerning this issue have been presented. In the case of other HA-based composites, Klc was in the range of 2.0 – 3.9 MPa m1/2, depending upon the reinforcement used. An undoubted advantage of the composite approach is an increase in toughness and strength of HA ceramics [2,3]. Tough and strong ceramics are produced with sintering process. ‘‘Sintering’’ is the term used to describe the consolidation of the product during firing. Consolidation implies that within the product, particles have joined together into strong aggregate. The term sintering is often interpreted to imply that shrinkage and densification have occurred. Sintered HA implants are observed to develop cracks [6]. Thus, an optimization of the strength and microstructure of HA implants by a suitable choice of sintering parameters are needed. It has been reported that the critical sintering temperature for calcium phosphates with
143
Ca/P ratio between 1.5 and 1.7 is 1300 jC [7]. Sintering of HA is complicated by the fact that HA is a hydrated phase that decomposes to anhydrous calcium phosphates such as TCP at f 1200– 1450 jC. Decomposition results from dehydroxylation beyond a critical point. For temperatures below the critical point (1300 jC), the HA crystal structure is retained despite dehydroxylation, and the HA rehydrates on cooling. If the critical point is exceeded, complete and irreversible dehydroxylation occurs, resulting in collapse of the HA structure and decomposition. Significant reversible dehydroxylation generally occurs above f 800 jC. After the critical point, a-TCP and h-TCP are often formed. In particular, the molecular volume increase that occurs in the hTCP ! a-TCP transformation seems to be the most deleterious phenomenon for mechanical properties [8].
2. Materials and methods 2.1. Production of HA The HA material used in this study was derived from freshly extracted human teeth. The teeth were irrigated with tap water and soaked in a 1% concen-
Fig. 1. X-ray diffraction diagram of the dentine sample.
144
G. Go¨ller, F.N. Oktar / Materials Letters 56 (2002) 142–147
tration of antiseptic solution to prevent bad odor and contamination of various infectious diseases [9]. Subsequently, the teeth were reirrigated and deproteinized in an alkali solution (1% concentration of sodium hypochloride). Thereafter, samples were reirrigated with tap water again. All samples were calcined at 5 jC min 1 to 850 jC and kept for 5 – 6 h. It was observed that at high temperatures, the dentine material and enamel material separated easily and that nearly 60% of the material was dentine and 40% enamel. Enamel and dentine particles were ground with a blade grinder for 30 s. Samples were poured into sieves placed from coarsest to finest in series. Dentine HA particles between 106 and 150 Am were used [10]. X-ray diffraction analysis of dentine particles has been carried out by Rigaku Rint-100 X-ray Diffractometer. Diffraction results indicated that the dentine material has 100% crystalline structure (Fig. 1). Synthetic HA powder ‘for bioceramics’ with a particle size in the range of 2 – 20 Am (MERCK 2196; Merck, Darmstadt, Germany) was also used for comparison. 2.2. Preparation of samples A die had been designed to prepare samples with a diameter of 12 mm according to a British Standard for
compression tests (no. 7253). The HA powder was weighed to total 3.25 g portions. The prepared portions were pressed at 350 MPa between hardened steel dies (die height 5 mm). In this study, any lubricant like polyvinylalcohol was used. All material were kept in normal atmospheric conditions and kept slightly humid for of a couple days (since dried HA powders would not compact and hold in one piece in the absence of humidity). Pressed samples were sintered in an open atmospheric oven at temperatures 1000, 1100, 1200, and 1300 jC (with the heating rate of + 4 jC min 1) for 4 h. The density of sintered samples was determined by the Archimedes method. Vickers hardness values were obtained using a Shimadzu microhardness testing machine with 200 g load. A universal test machine (Instron 1186) was used for compression tests.
3. Results and discussion The densification of HA as a function of sintering temperature follows a sigmoidal correlation with the attainment of plateau densification levels at 1100– 1300 jC. The plateau temperature, and associated limiting densification level, depend predominantly on the surface area of the HA powder and to a lesser extent on the heating rate and Ca/P ratio [11]. Accord-
Fig. 2. Sinterability of HA at various temperatures.
G. Go¨ller, F.N. Oktar / Materials Letters 56 (2002) 142–147
145
Fig. 3. Variation in compressive strength with sintering temperature.
ing to Suchanek et al. [3], full density can be achieved after sintering at 1400 jC for 3 h. The density of samples for different sintering temperatures obtained in this research can be seen in Fig. 2. The density of HA samples without sintering is qavr = 2.889 g/cm3. It is seen that sintering increases the density of the HA samples. A decrease in density occurring between 1000 and 1100 jC could be attributed to the initiation of h-TCP to a-TCP at that temperature range.
Although a slight increase is observed at 1200 and 1300 jC, the density of samples at 1000 jC is never exceeded. This is probably due to the relatively higher content of h-TCP phase at this temperature. The compression strength and the Vickers hardness values for sintered samples are shown in Figs. 3 and 4, respectively. There is an important increase of compression strength between 1100 and 1200 jC whereas compression strength values is about the same level at
Fig. 4. Variation in Vickers hardness with sintering temperature.
146
G. Go¨ller, F.N. Oktar / Materials Letters 56 (2002) 142–147
1200 – 1300 jC, as seen in Fig. 3. Therefore, it is meaningful to assume that the temperature between 1200 and 1300 jC is the optimum sintering temperature for maximum strength for biologically derived HA. In this study, it is observed that slightly lower results in strength and higher values in density and hardness are obtained compared to studies carried out with the HA powders derived from bone (Table 1). The mechanical test results suggest that the decomposition of biologically derived HA does not start at this temperature range. It is more stable than the synthetic powders commonly used in literature. Three recent investigations have correlated mechanical properties with sintering temperature for HA sintered parts at this range of temperatures. It is stated in literature that the microhardness plateau is similar to the sintering sigmoidal curve, while fracture toughness peaked at 1200 jC and dropped by f 20% at 1250 jC [12]. It is also stated that microhardness and fracture toughness both peaked at the sintering-plateau level, showing a significant decrease about 20– 40% at 1250 – 1300 jC [13]. Moreover, Go¨kbayrak [6] also used biologically derived HA with precisely prepared samples for the BS standard. The maximum strength value he had obtained is given in Table 1. As seen, the compression strength values obtained in this research are close, but density and hardness values are higher than that of Go¨kbayrak data. Wang and Chaki [7] had achieved the maximum density as 2.46 g cm 3 at 1100 jC for synthetic HA. They achieved the maximum hardness also at 1100 jC. It was seen that production of dentine HA samples were very efficient. There was no problem during material preparation whereas samples originated from synthetic HA (2– 20 Am) produced defects [6]. The highest hardness values are obtained at 1300 jC sintering temperature where density and compression strength values are also improved. This result suggests that powders used in this study show better
consolidation properties at higher temperatures. It would be better to perform a chemical analysis to obtain the decomposition rates of the materials at these temperatures for further studies. Ruys et al. [11] found that the apparent density decreased significantly above 1300 jC. The %HA – temperature curve indicated that the critical temperature above which decomposition occurred was 1350 jC. Therefore, 1350 jC is a reasonable value to assign for a nominal value of the decomposition temperature. Additionally, Ruys et al. stated that microstructural development is not dependent simply on sintering kinetics. Furthermore, dehydroxylation effects play an important role. A combination of the sintering and dehydroxylation effects points at two regions of importance in terms of the temperature – strength correlation; first, the attainment of the closed porosity densification level of f 95% at f 1150– 1200 jC and, second, decomposition of HA above 1350 jC. Optimum sintering temperature for synthetic HA is 1100– 1150 jC; above this temperature, it is possible to see phase transformations [14]. The overall results of sintering HA for 4 h at temperature range 1000 –1300 jC are shown in Table 1. Table 1 presents that sintering at 1300 jC makes possible to reach maximum hardness, compressive strength and density for biologically derived HA powders.
4. Conclusions Results indicated that it is possible to obtain dense HA with high strength, hardness and high density by sintering at 1300 jC for 4 h economically, which is suitable to be used as a graft material or tooth substitute. Modifications in sintering conditions and addition of reinforcing materials should be tested by the further studies.
Table 1 Results of sintering at various temperatures Temperature (jC)
qavr (g cm 3), Go¨kbayrak
ravr (MPa), Go¨kbayrak
Hardnessavr (HV), Go¨kbayrak
qavr (g cm 3), Gu¨ltekin
ravr (MPa), Gu¨ltekin
Hardnessavr (HV), Gu¨ltekin
1000 1100 1200 1300
2.46 2.34 2.59 2.48
48.17 22.16 75.20 65.01
85.37 79.20 148.50 130.68
2.992 2.945 2.958 2.961
9.83 14.48 56.30 56.77
2.58 163.1 229.8 751.9
G. Go¨ller, F.N. Oktar / Materials Letters 56 (2002) 142–147
Acknowledgements The authors would like to thank Prof. Dr. Selim Kusefoglu for his contribution during the sintering experiments, and Hakan Ulutas from Devotrans for compression tests.
References [1] J.A. Koempel, B.S. Patt, K. O’Grady, J. Wozney, D.M. Toriumi, J. Biomed. Mater. Res. 41 (1998) 359. [2] W. Suchanek, M. Yashima, M. Kakihana, M. Yoshimura, J. Am. Ceram. Soc. 80 (1997) 2805. [3] W. Suchanek, M. Yashima, M. Kakihana, M. Yoshimura, Biomaterials 17 (1996) 1715. [4] M. Knepper, S. Moricca, B.K. Milthorpe, Biomaterials 18 (1997) 1523. [5] G.D. With, A.J. Corbijn, J. Mater. Sci. 24 (1989) 3411. [6] H. Go¨kbayrak, Production of hydroxyapatite ceramics, MS thesis from the Mechanical Engineering Department of Bog˘azicßi University, 1993.
147
[7] P.E. Wang, T.K. Chaki, J. Mater. Sci.: Mater. Med. 4 (1993) 150. [8] A. Tampieri, G. Celotti, F. Szontagh, E. Landi, J. Mater. Sci.: Mater. Med. 8 (1997) 29. [9] S.V. Joshi, M.P. Srivastava, A. Pal, S. Pal, J. Mater. Sci.: Mater. Med. 4 (1993) 251. [10] F.N. Oktar, K. Kesenci˙, E. Pi˙ßs ki˙n, Int. J. Artif. Cells Blood Substitutes Immobilization Biotechnol. 27 (1999) 367. [11] A.J. Ruys, M. Wei, C.C. Sorrel, M.R. Dickson, A. Brandwood, B.K. Milthorpe, Biomaterials 16 (1995) 409. [12] J.D. Santos, J.C. Knowles, F.J. Monteiro, G.W. Hastings, Development of a glass reinforced hydroxyapatite with enhanced mechanical properties: physical characterisation and in vitro studies, in: T. Yamamuro, T. Kokubo, T. Nakamura (Eds.), Bioceramics, vol. 5, Kobunshi Kankokai, Kyoto, Japan, 1992, pp. 35 – 41. [13] S. Best, W. Bonfield, C. Doyle, Optimisation of toughness in dense hydroxyapatite ceramics, in: G. Heimke (Ed.), Bioceramics, vol. 2, Proceedings of the 2nd International Symposium on Ceramics in Medicine, DKG, Cologne, 1990, pp. 57 – 64. [14] T. Kitsugi, T. Yamamuro, T. Nakamura, S. Kotani, T. Kokubo, H. Takeuchi, Biomaterials 14 (1993) 216.