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Printing of Titanium implant prototype
Materials and Design 31 (2010) S101–S105
Contents lists available at ScienceDirect
Materials and Design journal homepage: www.elsevier.com/locate/matdes
Printing of Titanium implant prototype Florencia Edith Wiria a,*, John Yong Ming Shyan a, Poon Nian Lim a, Francis Goh Chung Wen a, Jin Fei Yeo b, Tong Cao b a b
Forming Technology Group, Singapore Institute of Manufacturing Technology, 71 Nanyang Drive, Singapore 638075, Singapore Department of Oral and Maxillofacial Surgery, Faculty of Dentistry, National University of Singapore, 5 Lower Kent Ridge Road, Singapore 119074, Singapore
a r t i c l e
i n f o
Article history: Received 14 July 2009 Accepted 31 December 2009 Available online 7 January 2010 Keywords: Titanium 3-Dimensional printing Scaffold Low modulus
a b s t r a c t Dental implant plays an important role as a conduit of force and stress to flow from the tooth to the related bone. In the load sharing between an implant and its related bone, the amount of stress carried by each of them directly related to their stiffness or modulus. Hence, it is a crucial issue for the implant to have matching mechanical properties, in particular modulus, between the implant and its related bone. Titanium is a metallic material that has good biocompatibility and corrosion resistance. Whilst the modulus of the bulk material is still higher than that of bone, it is the lowest among all other commonly used metallic implant materials, such as stainless steel or cobalt alloy. Hence it is potential to further reduce the modulus of pure Titanium by engineering its processing method to obtain porous structure. In this project, porous Titanium implant prototype is fabricated using 3-dimensional printing. This technique allows the flexibility of design customization, which is beneficial for implant fabrication as tailoring of implant size and shape helps to ensure the implant would fit nicely to the patient. The fabricated Titanium prototype had a modulus of 4.8—13.2 GPa, which is in the range of natural bone modulus. The compressive strength achieved was between 167 to 455 MPa. Subsequent cell culture study indicated that the porous Titanium prototype had good biocompatibility and is suitable for bone cell attachment and proliferation. Ó 2010 Elsevier Ltd. All rights reserved.
1. Introduction Dental-related diseases are one of the most common universal diseases across the world. There are various factors contributing to dental diseases, such as tooth corrosion due to frequent digestion of processed acidic food and drinks, poor personal oral hygiene, accidents or old age. In severe cases, tooth extraction may occur. After an extraction of a tooth, the related bone structure at the alveolar ridge is likely to experience disuse atrophy if the site is left vacant without any corrections [1]. The bone resorption is commonly due to the lack of stimulation from the chewing action of the tooth. It will finally cause the bone volume to be reduced. With time, the bone loss becomes more severe and eventually bone will be lost. Dental implants have become a common treatment alternative for the replacement of missing tooth in edentulous patients. A dental implant serves several functions. First, it serves as a base for the newly implanted prosthetic crown. Second, a dental implant re-
* Corresponding author. Tel.: +65 6793 8274; fax: +65 6971 6377. E-mail addresses: fl[email protected], fl[email protected] (F.E. Wiria). 0261-3069/$ - see front matter Ó 2010 Elsevier Ltd. All rights reserved. doi:10.1016/j.matdes.2009.12.050
places the root structure and serves as a conduit for stress, caused by the chewing action, to flow down from the newly implanted crown to the jaw bone. As the jaw bone continues to be active, this motion preserves prevents the bone related to tooth extraction from deteriorating. In addition, a dental implant also helps in restoring the tooth function, preserving teeth, improving a person’s esthetics and improving speech problems. Dental implant has been reported to have a considerable amount of market size, with an estimated size US243.3 million with a growth of 16% [2]. For implant with load bearing function, the load is ideally shared between the bone and the implant. The amount of stress carried by each of them is directly related to their stiffness or elastic modulus [3]. If the implant has an elastic modulus far greater than that of the bone, there is a concern that the higher elastic modulus of the implant would take the majority of the stress loaded and thus preventing the bone from being loaded properly [4]. This phenomena is called stress shielding, with the adverse effect of implant loosening in the long term [5]. Therefore, it is important for the dental implant and the related bone to have similar elastic modulus. Currently available dental implants are mostly made of Titanium, Zirconia and Alumina materials [6]. Alumina and zirconia are used due to their high strength, however dental ceramics are brittle and commonly have low fracture resistance and thus limit
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F.E. Wiria et al. / Materials and Design 31 (2010) S101–S105
the possibility of manufacturing them [7]. In recent years, there has also been research going on the use of poly-lactic-acid and its copolymers for dental implant [8]. However, these polymers in general have relatively much lower strength and modulus values as compared to bone [9]. Titanium is a metallic material with good biocompatibility and corrosion resistance. It has been widely used for implants, such as prosthetic joints, dense dental fixtures, screws and heart valves. The modulus of bulk pure Titanium is 102–105 GPa [10]. Whilst the modulus of the bulk material is still higher than that of bone, it is the lowest among all other commonly used metallic implant materials, such as stainless steel or cobalt–chromium alloy [3]. Table 1 summarizes the mechanical properties of natural jaw bone and several bulk biomaterials that are commonly used for implants. As seen in Table 1, bulk metal and ceramics have properties in both strength and modulus higher to that of bone. Therefore, considering its favorable properties, pure Titanium has the best potential to be engineered to have its modulus matched to that of bone. By introducing porosity to the bulk pure Titanium through engineering its processing method, its modulus could be reduced to match that of bone. Apart from further reduction in the modulus of the bulk material, the porosity in the dental implant provides channels for the cells to grow inside the porous implant [11]. A layer-by-layer fabrication method is explored to produce the implant green part for this project. This fabrication method commonly requires the object to be first drawn in a computer model representation. As such, the model could be easily modified in terms of dimension and shape according to the needs this approach of fabrication is beneficial for implant fabrication, as the implant could be customized accordingly [12]. Several layer-bylayer fabrication techniques have been researched for the purpose of fabricating customized implants [13–17]. A 3D printer (3DP) is a layer-by-layer manufacturing technique based on inkjet printing deposition concept. A 3DP is mainly equipped with a powder supply chamber, a part building chamber, a roller and an inkjet printer depositor. It works in the following manner, as can be seen from Fig. 1. The powder supply chamber moves up one layer thickness in height to supply a layer of powder material, which is then transferred onto the printing chamber with the help of the roller. The ink depositor mechanism then moves to selectively drop the liquid adhesive binder onto the cross-sectional area of the sliced object. This action produces a layer of bonded powder material at the selected regions. The part building chamber then moves down one layer thickness in height after the printing is done followed by upward movement of the powder supply chamber to supply a fresh layer of powder. These steps are repeated until the part is finally formed. The unprinted powder becomes a natural support as the object is built and is eventually cleaned from the final part. The objectives of this project are to firstly, fabricate low modulus porous Titanium implant via printing method. Secondly, to
Table 1 Mechanical properties of biomaterials. Material
Strength (MPa)
Modulus (GPa)
Natural jaw bone PLGA Stainless steel 316L Co–Cr–Mo Ti–6Al–4V Commercially pure Titanium Alumina Yttria-stabilized zirconia
130–180 2.82 170–750 275–1585 895–930 240–550 400 900–1400
3–20 2 200 200–230 110–114 102–105 350 210
Ink depositor Printed cross section area
roller
Part building chamber
Powder supply chamber
Fig. 1. 3D printing process.
demonstrate the suitability of the fabricated porous implant prototype as an environment for in-vitro cell culture growth. 2. Materials and method 2.1. Fabrication of dental implant prototypes The Titanium implant is first designed in a computer-aided design (CAD) environment (Pro-Engineer Wildfire). Afterwards, the design is transferred into a .STL (stereolithography) file extension, where the 3D volumetric object is virtually sliced into very thin cross-sectional layers for fabrication process using a layer-by-layer manufacturing concept. The implant prototype is then transferred to a 3DP for fabrication. A 3DP (Z Corporation, model 310 Plus) is used to fabricate the green part of the Titanium implant prototype. The Commercially Pure (CP) Titanium powder ( 325 mesh) used is supplied by Alfa Aesar (99.5% purity). Poly(vinyl alcohol) (PVA) powder (75 lm) is selected to be used as the binder powder (Nippon Gohsei, NH18S). These powders are mixed using a tumbler mixer (Bioengineering Inversina™). As the 3DP is commercialized equipment that is initially specialized for fabricating industrial components, their propriety materials are plaster based. With the use of a non propriety material, selected functions and parameters are modified to accommodate the printing of the metallic powder. Debinding and sintering of the printed Titanium prototype green parts is performed in Argon atmosphere at the rate of 20 l/ min (CM Tube High Temperature Furnace). In this process, the green parts are firstly debound from the PVA binder and subsequently sintered. Table 2 provides the summary of the sintering condition. Mechanical strength of the Titanium prototypes is characterized by compression testing, based on ASM Standards of Axial Compression Testing [18]. Five sintered samples (n = 5) for each of the three different sintering parameters are tested using universal testing machine (INSTRONÒ) at a strain rate of 0.5 mm/min. The specimens are cylindrical in shape, with a width to height ratio of 1:3. Microstructural observation of the Titanium parts is conducted by cutting the samples in their cross-section using a Struers™ diamond cutter wheel. The sample is mounted in epoxy resin and cured overnight. The observation is conducted using a scanning electron microscope (JEOL SEM 56000LV) using a voltage of 20 kV.
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F.E. Wiria et al. / Materials and Design 31 (2010) S101–S105 Table 2 Sintering conditions of Titanium green parts. Sintering temperature (°C)
Holding time (min)
Pressure (atm)
Heating rate (°C/ min)
1250 1300 1350
120
1
5
2.2. Cell culture study An in-vitro cell culture study is carried out by seeding cells onto the porous Titanium structures. A cytotoxicity study is carried out to investigate if the fabricated porous Titanium implant prototypes are harmful to cells. These implants function as scaffold to support the newly seeded cells and guide their growth into the intended shape. Agarose gel cylinders with the same dimension as Titanium implants were used as negative control for cytotoxicity test; the same material with encapsulation of phenol were used as positive control. The agarose gel cylinders were prepared from 1.5% (w/v) agarose which was melted at 120 °C for 20 min. L-929 fibroblast cells (CCL-1, ATCC) are seeded at 5104 cell/ cm2 in a 6-well plate and incubated for 12 h. The sterilized scaffolds, negative control cylinders and positive control cylinders are subsequently placed in the center of L-929 cell-culture wells in triplicates and further incubated for another period of 48 h with 1 ml of fresh media. After removing the scaffolds and reference cylinders, cell viability is quantitatively analyzed with CellTiter 96 AQueous Non-Radioactive Cell Proliferation Assay (MTS) kit (Promega Corp, USA). Colorimetric analysis was subsequently performed by reading 490 nm absorbance with an Infinite 200 microplate reader (Tecan Trading AG, Switzerland). Cytotoxicity of the Titanium scaffold is reported by percentage of cell viability. Seeding of h-FOB 1.19 osteoblast cells (CRL-11372, ATCC) are also conducted to check if the Titanium implants are suitable for other types of cells. After 14 days of continuous culture, fluorescein diacetate (FDA) (Sigma) and propidium iodide (PI) (Sigma) staining is performed to view the cell attachment directly under fluorescent microscope. The implants are placed in 1 ml of 2% of FDA/PI (1:1) for 5mins and washed with PBS completely before viewing. FDA stains the cytoplasm of living cells and PI stains the nucleus of dead cells.
3. Results and discussion 3.1. Dental implant prototype The Titanium implant prototype fabricated is shown in Fig. 2. Fig. 2a shows the printed green printed part and Fig. 2b shows a typical sintered part. Measurement of the diameter and height of the sintered Titanium prototypes shows that the prototypes experience 27–28% shrinkage in diameter and 21–26% shrinkage in height, when sintered at temperature 1250, 1300 and 1350 °C. Table 3 shows the shrinkage experienced by the sintered Titanium parts. With the increase in the sintering temperature, the shrinkage experienced by the parts also increases. The shrinkage in the height is observed to be relatively higher than the shrinkage in the diameter. This is caused by the placement of the green part during the sintering process. The Titanium green part consists of three columns, which is linked by the circular barriers at the top and bottom. During sintering, the parts are placed with the height in the z-direction, in the manner that the prototypes were standing, like shown in Fig. 2a. In most cases of sintering, capillary force is usually known as the main driving force
Fig. 2. Titanium implant prototype: (a) green printed part and (b) sintered part at 1250 °C.
leading to the rearrangement of powder particles in sintering process [19]. The capillary force is generally considered to cause isotropic shrinkage. However, there is also the gravitational force acting on the component during sintering, and the gravitational force is not uniform in x-, y- and z-directions [20]. The largest would be in the z-direction, which is the direction of the columns. As the PVA particles in the green parts are burnt during the debinding process, they left behind void volume. The Titanium particles are then ‘‘pulled down” due to this gravitational force to fill in the void volume. As the time taken to reach the sintering temperature was longer as the sintering temperature was increased, the more likely that the gravitational force acting on the columns to push them down as they become denser due to the sintering of the Titanium part. Therefore, it causes the parts more compact in the z-direction as compared to the lateral direction. During the sintering process, particle bonding is achieved through solid state diffusion process, which causes necking growth, firstly at the point of contact between adjacent particles. This is then continued by neck formation at the boundary of the particles [21]. Fig. 3 shows the pore structure of the Titanium part sintered at different temperatures. It can be seen that the portions of dense areas increase with the increase in sintering temperature. Parts sintered with lower sintering temperature gives more open pores with higher degree of interconnectivity to form porous network. However, more of these open pores are closed with the increase of the sintering temperature, therefore causing lesser degree of pore interconnectivity. Table 2 shows the elastic modulus and compressive strength of the Titanium prototypes when sintered at various temperatures. It can be seen that the elastic modulus and compressive strength of the porous Titanium structures varies with the sintering temperature. The compressive properties increase with the increase in the sintering temperature. As known from the previous observation, the increase in the sintering temperature decreases the degree of porosity and pore interconnectivity, resulting in a part with higher dense volume. The current observation then illustrates that by increasing the sintering temperature, the mechanical properties are also increased, mainly due to the increase in the dense volume of the part. It is also seen the elastic modulus of the Titanium structures obtained by varying the sintering profiles falls within the elastic modulus range of the natural bone, which is 3–20 GPa [22]. This shows the capability of the process to tailor the Titanium parts
Table 3 Shrinkage experienced by the sintered Titanium parts. Sintering temperature (°C)
Shrinkage in diameter (%)
Shrinkage in height (%)
1250 1300 1350
27.8 ± 6.1 28.4 ± 3.2 28.7 ± 1.9
21.2 ± 3.4 23.4 ± 1.9 26.1 ± 1.7
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F.E. Wiria et al. / Materials and Design 31 (2010) S101–S105 Table 4 Mechanical properties of the Titanium prototypes. Sintering temperature (°C)
Elastic modulus (GPa)
Compressive strength (MPa)
1250 1300 1350 Natural bone
4.8 10.8 13.2 3–20
166.9 408.2 455.1 130–180
Fig. 4. Dimensions measured for repeatability test of the process chain.
outer and inner diameter. The prototype is designed to have a final dimension of 5-mm in both diameter and height. Measurements for diameters 1 and 2 are (4.89 ± 0.07) and (5.42 ± 0.15) mm, respectively. Thicknesses 1 and 2 are (1.31 ± 0.06) and (1.33 ± 0.06) mm, respectively, and the measurement of the height is (5.11 ± 0.14) mm. The measurement results show that the average dimension of the prototypes is very close to the actual desired dimension. The final prototypes have a maximum deviation of 4.6%, which indicates good repeatability of the fabrication process.
3.2. In-vitro study of Titanium scaffold
Fig. 3. Micrographs of Titanium parts (magnified 200) sintered at: (a) 1250 °C, (b) 1300 °C and (c) 1350 °C.
such that the properties can be tailored. By deliberately engineering the Titanium prototypes porous it helps to reduce the modulus of the bulk pure Titanium and meet the modulus of natural bone. This could help preventing adverse effect of bone loosening due to mismatch in the elastic modulus between implant and the affected bone (see Table 4). The process stability is carried out by measuring the dimensions of the Titanium implant prototypes (n = 20) using a vernier caliper. The dimensions measured are the height, diameter and thickness of the dental cage circumference, all of which are illustrated in Fig. 4. The diameter is referred to be the outer diameter of the prototype and the thickness is the nominal difference between the
In this in-vitro study, the positive control is a material that would give a 100% positive result. Hence for this cytotoxicity test the positive control is referred to sample that gives positive toxic result, which means that the sample is toxic for cells. On the contrary, the negative control is a sample that is toxic to cells. The bar chart comparison of cytotoxicity between implant and reference materials is presented in Fig. 5. Very minimum difference in L-929 cell viability is observed between the group of Titanium scaffolds and the negative control (P = 0.7667). This indicates that the toxicity level of the Titanium scaffold group is very near to that of the negative control. On the contrary, the cell viability of Titanium scaffold group is very significantly higher than that of positive control group (P < 0.001). Therefore, it is concluded that the Titanium scaffold displays an extremely low cytotoxic effect on standard cells. After being continuously cultured for 14 days, the surface of the Titanium scaffolds is fully covered with living cells which are stained by FDA, as seen by the green color in Fig. 6. Among these cells, only a minority of dead cells are observed. These dead cells are stained by PI, as seen by the red color in Fig. 6. This second cell culture study of seeding osteoblast (bone cells) onto the Titanium scaffolds has shown that the bone cells could attach on the scaffolds and that they could proliferate well within 14 days with minimum amount of dead cells. This shows that the Titanium implant prototypes are able to provide suitable substrate for cells to attach and live.
F.E. Wiria et al. / Materials and Design 31 (2010) S101–S105
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In-vitro cell culture results have presented good and promising results. The fabricated porous Titanium dental implants are demonstrated to have very low toxicity effect. The implants have also shown that they can be used as suitable substrate for bone cells to live, grow and proliferate. In summary, as a bone-engineering material, the Titanium implant prototypes present good mechanical properties and biocompatibility by in-vitro investigation as stratum that supports cell adherence, cell proliferation and osteogenic differentiation, with little cytotoxicity effect. References
Fig. 5. Comparison of cytotoxicity between Titanium scaffolds and control cylinders.
Fig. 6. FDA and PI staining of cells on Titanium scaffolds 14 days after seeding.
4. Conclusions Commercially pure Titanium has been successfully investigated to fabricate porous dental implant prototype. Both the fabrication and cell culture results have been satisfactory. The dental implant prototype has been successfully fabricated using printing method and shown to have elastic modulus of 4.8–13.2 GPa. This elastic modulus is much lower than the modulus of the bulk commercially pure Titanium and is in the range of elastic modulus of natural bone [22]. The fabrication method also allowed variation of mechanical properties, as needed by the bone.
[1] Ihde S. Four-dimensional considerations of bone morphology and mechanics. In: Ihde S, editor. Principles of BOI: clinical scientific and practical guidelines to 4-D dental implantology. Berlin: Springer; 2005. p. 103–43. [2] Callan D. Dental implants. In: Vernino AR, Gray J, Hughes E, editors. The periodontic syllabus. Philadelphia: Lippincott Williams & Wilkins; 2007. [3] Ramakrishna S, Mayer J, Wintermantel E, Leong KW. Biomedical applications of polymer-composite materials: a review. Compos Sci Technol 2001;61(9):1189–224. [4] Hench LL. Bioceramics: from concept to clinic. J Am Ceram Soc 1991;74(7): 1487–510. [5] Hench LL. Bioceramics. J Am Ceram Soc 1998;81(7):1705–28. [6] Kohal RJ, Att W, Baechle M, Butz F. Ceramic abutments and ceramic oral implants. An update. Periodontology 2000 2008;47:224–43. [7] Sundh A, Molin M, Sjögren G. Fracture resistance of yttrium oxide partiallystabilized zirconia all-ceramic bridges after veneering and mechanical fatigue testing. Dent Mater 2005;21(5):476–82. [8] Mastiholimath VS, Dandagi PM, Gadad AP, Patil MB, Manvi FV, Chandur VK. Formulation and evaluation of ornidazole dental implants for periodontitis. Indian J Pharm Sci 2006;68(1):68–71. [9] Martina M, Hutmacher DW. Biodegradable polymers applied in tissue engineering research: a review. Polym Int 2007;56(2):145–57. [10] Lütjering G, Williams JC. Titanium. Heidelberg: Springer-Verlag; 2007. [11] Olivier V, Faucheux N, Hardouin P. Biomaterial challenges and approaches to stem cell use in bone reconstructive surgery. Drug Discov Today 2004;9(18):803–11. [12] Chua CK, Leong KF, Lim CS. Rapid prototyping: principles and applications. 2nd ed. Singapore: World Scientific; 2003. [13] Wiria FE, Leong KF, Chua CK, Liu Y. Poly-e-caprolactone/hydroxyapatite for tissue engineering scaffold fabrication via selective laser sintering. Acta Biomater 2007;3(1):1–12. [14] Yang SF, Leong KF, Du ZH, Chua CK. The design of scaffolds for use in tissue engineering. Part II. Rapid prototyping techniques. Tissue Eng 2002;8(1):1–11. [15] Yeong WY, Chua CK, Leong KF, Chandrasekaran M. Rapid prototyping in tissue engineering: challenges and potential. Trends Biotechnol 2004;22(12):643–52. [16] Wiria FE, Chua CK, Leong KF, Quah ZY, Chandrasekaran M, Lee MW. Improved biocomposite development of poly(vinyl alcohol)/hydroxyapatite for tissue engineering scaffolds fabrication using selective laser sintering. J Mater Sci: Mater Med 2008;19(3):989–96. [17] Simpson RL, Wiria FE, Amis AA, Chua CK, Leong KF, Hansen UN, Chandrasekaran M, Lee MW. Development of a 95/5 poly(L-lactide-coglycolide)/hydroxyapatite & b-tricalcium phosphate scaffold as bone replacement material via selective laser sintering. J Biomed Mater Res: Part B – Appl Biomater 2008;84B:17–25. [18] Papirno R. ASM standard for axial compression testing. Ohio: ASM International. [19] German RM. Liquid phase sintering. New York: Plenum Press; 1985. [20] Zhu HH, Lu L, Fuh JYH. Study on shrinkage behaviour of direct laser sintering metallic powder. Proc Inst Mech Eng, Part B – J Eng Manuf 2006;220(2):183–90. [21] Rosenzweig N, Narkis M. Dimensional variations of two spherical polymeric particles during sintering. Polym Eng Sci 1981;21(10):582–5. [22] Krishna BV, Bose S, Bandyopadhyay A. Low stiffness porous Ti structures for load-bearing implants. Acta Biomater 2007;3(6):997–1006.
Contents lists available at ScienceDirect
Materials and Design journal homepage: www.elsevier.com/locate/matdes
Printing of Titanium implant prototype Florencia Edith Wiria a,*, John Yong Ming Shyan a, Poon Nian Lim a, Francis Goh Chung Wen a, Jin Fei Yeo b, Tong Cao b a b
Forming Technology Group, Singapore Institute of Manufacturing Technology, 71 Nanyang Drive, Singapore 638075, Singapore Department of Oral and Maxillofacial Surgery, Faculty of Dentistry, National University of Singapore, 5 Lower Kent Ridge Road, Singapore 119074, Singapore
a r t i c l e
i n f o
Article history: Received 14 July 2009 Accepted 31 December 2009 Available online 7 January 2010 Keywords: Titanium 3-Dimensional printing Scaffold Low modulus
a b s t r a c t Dental implant plays an important role as a conduit of force and stress to flow from the tooth to the related bone. In the load sharing between an implant and its related bone, the amount of stress carried by each of them directly related to their stiffness or modulus. Hence, it is a crucial issue for the implant to have matching mechanical properties, in particular modulus, between the implant and its related bone. Titanium is a metallic material that has good biocompatibility and corrosion resistance. Whilst the modulus of the bulk material is still higher than that of bone, it is the lowest among all other commonly used metallic implant materials, such as stainless steel or cobalt alloy. Hence it is potential to further reduce the modulus of pure Titanium by engineering its processing method to obtain porous structure. In this project, porous Titanium implant prototype is fabricated using 3-dimensional printing. This technique allows the flexibility of design customization, which is beneficial for implant fabrication as tailoring of implant size and shape helps to ensure the implant would fit nicely to the patient. The fabricated Titanium prototype had a modulus of 4.8—13.2 GPa, which is in the range of natural bone modulus. The compressive strength achieved was between 167 to 455 MPa. Subsequent cell culture study indicated that the porous Titanium prototype had good biocompatibility and is suitable for bone cell attachment and proliferation. Ó 2010 Elsevier Ltd. All rights reserved.
1. Introduction Dental-related diseases are one of the most common universal diseases across the world. There are various factors contributing to dental diseases, such as tooth corrosion due to frequent digestion of processed acidic food and drinks, poor personal oral hygiene, accidents or old age. In severe cases, tooth extraction may occur. After an extraction of a tooth, the related bone structure at the alveolar ridge is likely to experience disuse atrophy if the site is left vacant without any corrections [1]. The bone resorption is commonly due to the lack of stimulation from the chewing action of the tooth. It will finally cause the bone volume to be reduced. With time, the bone loss becomes more severe and eventually bone will be lost. Dental implants have become a common treatment alternative for the replacement of missing tooth in edentulous patients. A dental implant serves several functions. First, it serves as a base for the newly implanted prosthetic crown. Second, a dental implant re-
* Corresponding author. Tel.: +65 6793 8274; fax: +65 6971 6377. E-mail addresses: fl[email protected], fl[email protected] (F.E. Wiria). 0261-3069/$ - see front matter Ó 2010 Elsevier Ltd. All rights reserved. doi:10.1016/j.matdes.2009.12.050
places the root structure and serves as a conduit for stress, caused by the chewing action, to flow down from the newly implanted crown to the jaw bone. As the jaw bone continues to be active, this motion preserves prevents the bone related to tooth extraction from deteriorating. In addition, a dental implant also helps in restoring the tooth function, preserving teeth, improving a person’s esthetics and improving speech problems. Dental implant has been reported to have a considerable amount of market size, with an estimated size US243.3 million with a growth of 16% [2]. For implant with load bearing function, the load is ideally shared between the bone and the implant. The amount of stress carried by each of them is directly related to their stiffness or elastic modulus [3]. If the implant has an elastic modulus far greater than that of the bone, there is a concern that the higher elastic modulus of the implant would take the majority of the stress loaded and thus preventing the bone from being loaded properly [4]. This phenomena is called stress shielding, with the adverse effect of implant loosening in the long term [5]. Therefore, it is important for the dental implant and the related bone to have similar elastic modulus. Currently available dental implants are mostly made of Titanium, Zirconia and Alumina materials [6]. Alumina and zirconia are used due to their high strength, however dental ceramics are brittle and commonly have low fracture resistance and thus limit
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the possibility of manufacturing them [7]. In recent years, there has also been research going on the use of poly-lactic-acid and its copolymers for dental implant [8]. However, these polymers in general have relatively much lower strength and modulus values as compared to bone [9]. Titanium is a metallic material with good biocompatibility and corrosion resistance. It has been widely used for implants, such as prosthetic joints, dense dental fixtures, screws and heart valves. The modulus of bulk pure Titanium is 102–105 GPa [10]. Whilst the modulus of the bulk material is still higher than that of bone, it is the lowest among all other commonly used metallic implant materials, such as stainless steel or cobalt–chromium alloy [3]. Table 1 summarizes the mechanical properties of natural jaw bone and several bulk biomaterials that are commonly used for implants. As seen in Table 1, bulk metal and ceramics have properties in both strength and modulus higher to that of bone. Therefore, considering its favorable properties, pure Titanium has the best potential to be engineered to have its modulus matched to that of bone. By introducing porosity to the bulk pure Titanium through engineering its processing method, its modulus could be reduced to match that of bone. Apart from further reduction in the modulus of the bulk material, the porosity in the dental implant provides channels for the cells to grow inside the porous implant [11]. A layer-by-layer fabrication method is explored to produce the implant green part for this project. This fabrication method commonly requires the object to be first drawn in a computer model representation. As such, the model could be easily modified in terms of dimension and shape according to the needs this approach of fabrication is beneficial for implant fabrication, as the implant could be customized accordingly [12]. Several layer-bylayer fabrication techniques have been researched for the purpose of fabricating customized implants [13–17]. A 3D printer (3DP) is a layer-by-layer manufacturing technique based on inkjet printing deposition concept. A 3DP is mainly equipped with a powder supply chamber, a part building chamber, a roller and an inkjet printer depositor. It works in the following manner, as can be seen from Fig. 1. The powder supply chamber moves up one layer thickness in height to supply a layer of powder material, which is then transferred onto the printing chamber with the help of the roller. The ink depositor mechanism then moves to selectively drop the liquid adhesive binder onto the cross-sectional area of the sliced object. This action produces a layer of bonded powder material at the selected regions. The part building chamber then moves down one layer thickness in height after the printing is done followed by upward movement of the powder supply chamber to supply a fresh layer of powder. These steps are repeated until the part is finally formed. The unprinted powder becomes a natural support as the object is built and is eventually cleaned from the final part. The objectives of this project are to firstly, fabricate low modulus porous Titanium implant via printing method. Secondly, to
Table 1 Mechanical properties of biomaterials. Material
Strength (MPa)
Modulus (GPa)
Natural jaw bone PLGA Stainless steel 316L Co–Cr–Mo Ti–6Al–4V Commercially pure Titanium Alumina Yttria-stabilized zirconia
130–180 2.82 170–750 275–1585 895–930 240–550 400 900–1400
3–20 2 200 200–230 110–114 102–105 350 210
Ink depositor Printed cross section area
roller
Part building chamber
Powder supply chamber
Fig. 1. 3D printing process.
demonstrate the suitability of the fabricated porous implant prototype as an environment for in-vitro cell culture growth. 2. Materials and method 2.1. Fabrication of dental implant prototypes The Titanium implant is first designed in a computer-aided design (CAD) environment (Pro-Engineer Wildfire). Afterwards, the design is transferred into a .STL (stereolithography) file extension, where the 3D volumetric object is virtually sliced into very thin cross-sectional layers for fabrication process using a layer-by-layer manufacturing concept. The implant prototype is then transferred to a 3DP for fabrication. A 3DP (Z Corporation, model 310 Plus) is used to fabricate the green part of the Titanium implant prototype. The Commercially Pure (CP) Titanium powder ( 325 mesh) used is supplied by Alfa Aesar (99.5% purity). Poly(vinyl alcohol) (PVA) powder (75 lm) is selected to be used as the binder powder (Nippon Gohsei, NH18S). These powders are mixed using a tumbler mixer (Bioengineering Inversina™). As the 3DP is commercialized equipment that is initially specialized for fabricating industrial components, their propriety materials are plaster based. With the use of a non propriety material, selected functions and parameters are modified to accommodate the printing of the metallic powder. Debinding and sintering of the printed Titanium prototype green parts is performed in Argon atmosphere at the rate of 20 l/ min (CM Tube High Temperature Furnace). In this process, the green parts are firstly debound from the PVA binder and subsequently sintered. Table 2 provides the summary of the sintering condition. Mechanical strength of the Titanium prototypes is characterized by compression testing, based on ASM Standards of Axial Compression Testing [18]. Five sintered samples (n = 5) for each of the three different sintering parameters are tested using universal testing machine (INSTRONÒ) at a strain rate of 0.5 mm/min. The specimens are cylindrical in shape, with a width to height ratio of 1:3. Microstructural observation of the Titanium parts is conducted by cutting the samples in their cross-section using a Struers™ diamond cutter wheel. The sample is mounted in epoxy resin and cured overnight. The observation is conducted using a scanning electron microscope (JEOL SEM 56000LV) using a voltage of 20 kV.
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F.E. Wiria et al. / Materials and Design 31 (2010) S101–S105 Table 2 Sintering conditions of Titanium green parts. Sintering temperature (°C)
Holding time (min)
Pressure (atm)
Heating rate (°C/ min)
1250 1300 1350
120
1
5
2.2. Cell culture study An in-vitro cell culture study is carried out by seeding cells onto the porous Titanium structures. A cytotoxicity study is carried out to investigate if the fabricated porous Titanium implant prototypes are harmful to cells. These implants function as scaffold to support the newly seeded cells and guide their growth into the intended shape. Agarose gel cylinders with the same dimension as Titanium implants were used as negative control for cytotoxicity test; the same material with encapsulation of phenol were used as positive control. The agarose gel cylinders were prepared from 1.5% (w/v) agarose which was melted at 120 °C for 20 min. L-929 fibroblast cells (CCL-1, ATCC) are seeded at 5104 cell/ cm2 in a 6-well plate and incubated for 12 h. The sterilized scaffolds, negative control cylinders and positive control cylinders are subsequently placed in the center of L-929 cell-culture wells in triplicates and further incubated for another period of 48 h with 1 ml of fresh media. After removing the scaffolds and reference cylinders, cell viability is quantitatively analyzed with CellTiter 96 AQueous Non-Radioactive Cell Proliferation Assay (MTS) kit (Promega Corp, USA). Colorimetric analysis was subsequently performed by reading 490 nm absorbance with an Infinite 200 microplate reader (Tecan Trading AG, Switzerland). Cytotoxicity of the Titanium scaffold is reported by percentage of cell viability. Seeding of h-FOB 1.19 osteoblast cells (CRL-11372, ATCC) are also conducted to check if the Titanium implants are suitable for other types of cells. After 14 days of continuous culture, fluorescein diacetate (FDA) (Sigma) and propidium iodide (PI) (Sigma) staining is performed to view the cell attachment directly under fluorescent microscope. The implants are placed in 1 ml of 2% of FDA/PI (1:1) for 5mins and washed with PBS completely before viewing. FDA stains the cytoplasm of living cells and PI stains the nucleus of dead cells.
3. Results and discussion 3.1. Dental implant prototype The Titanium implant prototype fabricated is shown in Fig. 2. Fig. 2a shows the printed green printed part and Fig. 2b shows a typical sintered part. Measurement of the diameter and height of the sintered Titanium prototypes shows that the prototypes experience 27–28% shrinkage in diameter and 21–26% shrinkage in height, when sintered at temperature 1250, 1300 and 1350 °C. Table 3 shows the shrinkage experienced by the sintered Titanium parts. With the increase in the sintering temperature, the shrinkage experienced by the parts also increases. The shrinkage in the height is observed to be relatively higher than the shrinkage in the diameter. This is caused by the placement of the green part during the sintering process. The Titanium green part consists of three columns, which is linked by the circular barriers at the top and bottom. During sintering, the parts are placed with the height in the z-direction, in the manner that the prototypes were standing, like shown in Fig. 2a. In most cases of sintering, capillary force is usually known as the main driving force
Fig. 2. Titanium implant prototype: (a) green printed part and (b) sintered part at 1250 °C.
leading to the rearrangement of powder particles in sintering process [19]. The capillary force is generally considered to cause isotropic shrinkage. However, there is also the gravitational force acting on the component during sintering, and the gravitational force is not uniform in x-, y- and z-directions [20]. The largest would be in the z-direction, which is the direction of the columns. As the PVA particles in the green parts are burnt during the debinding process, they left behind void volume. The Titanium particles are then ‘‘pulled down” due to this gravitational force to fill in the void volume. As the time taken to reach the sintering temperature was longer as the sintering temperature was increased, the more likely that the gravitational force acting on the columns to push them down as they become denser due to the sintering of the Titanium part. Therefore, it causes the parts more compact in the z-direction as compared to the lateral direction. During the sintering process, particle bonding is achieved through solid state diffusion process, which causes necking growth, firstly at the point of contact between adjacent particles. This is then continued by neck formation at the boundary of the particles [21]. Fig. 3 shows the pore structure of the Titanium part sintered at different temperatures. It can be seen that the portions of dense areas increase with the increase in sintering temperature. Parts sintered with lower sintering temperature gives more open pores with higher degree of interconnectivity to form porous network. However, more of these open pores are closed with the increase of the sintering temperature, therefore causing lesser degree of pore interconnectivity. Table 2 shows the elastic modulus and compressive strength of the Titanium prototypes when sintered at various temperatures. It can be seen that the elastic modulus and compressive strength of the porous Titanium structures varies with the sintering temperature. The compressive properties increase with the increase in the sintering temperature. As known from the previous observation, the increase in the sintering temperature decreases the degree of porosity and pore interconnectivity, resulting in a part with higher dense volume. The current observation then illustrates that by increasing the sintering temperature, the mechanical properties are also increased, mainly due to the increase in the dense volume of the part. It is also seen the elastic modulus of the Titanium structures obtained by varying the sintering profiles falls within the elastic modulus range of the natural bone, which is 3–20 GPa [22]. This shows the capability of the process to tailor the Titanium parts
Table 3 Shrinkage experienced by the sintered Titanium parts. Sintering temperature (°C)
Shrinkage in diameter (%)
Shrinkage in height (%)
1250 1300 1350
27.8 ± 6.1 28.4 ± 3.2 28.7 ± 1.9
21.2 ± 3.4 23.4 ± 1.9 26.1 ± 1.7
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F.E. Wiria et al. / Materials and Design 31 (2010) S101–S105 Table 4 Mechanical properties of the Titanium prototypes. Sintering temperature (°C)
Elastic modulus (GPa)
Compressive strength (MPa)
1250 1300 1350 Natural bone
4.8 10.8 13.2 3–20
166.9 408.2 455.1 130–180
Fig. 4. Dimensions measured for repeatability test of the process chain.
outer and inner diameter. The prototype is designed to have a final dimension of 5-mm in both diameter and height. Measurements for diameters 1 and 2 are (4.89 ± 0.07) and (5.42 ± 0.15) mm, respectively. Thicknesses 1 and 2 are (1.31 ± 0.06) and (1.33 ± 0.06) mm, respectively, and the measurement of the height is (5.11 ± 0.14) mm. The measurement results show that the average dimension of the prototypes is very close to the actual desired dimension. The final prototypes have a maximum deviation of 4.6%, which indicates good repeatability of the fabrication process.
3.2. In-vitro study of Titanium scaffold
Fig. 3. Micrographs of Titanium parts (magnified 200) sintered at: (a) 1250 °C, (b) 1300 °C and (c) 1350 °C.
such that the properties can be tailored. By deliberately engineering the Titanium prototypes porous it helps to reduce the modulus of the bulk pure Titanium and meet the modulus of natural bone. This could help preventing adverse effect of bone loosening due to mismatch in the elastic modulus between implant and the affected bone (see Table 4). The process stability is carried out by measuring the dimensions of the Titanium implant prototypes (n = 20) using a vernier caliper. The dimensions measured are the height, diameter and thickness of the dental cage circumference, all of which are illustrated in Fig. 4. The diameter is referred to be the outer diameter of the prototype and the thickness is the nominal difference between the
In this in-vitro study, the positive control is a material that would give a 100% positive result. Hence for this cytotoxicity test the positive control is referred to sample that gives positive toxic result, which means that the sample is toxic for cells. On the contrary, the negative control is a sample that is toxic to cells. The bar chart comparison of cytotoxicity between implant and reference materials is presented in Fig. 5. Very minimum difference in L-929 cell viability is observed between the group of Titanium scaffolds and the negative control (P = 0.7667). This indicates that the toxicity level of the Titanium scaffold group is very near to that of the negative control. On the contrary, the cell viability of Titanium scaffold group is very significantly higher than that of positive control group (P < 0.001). Therefore, it is concluded that the Titanium scaffold displays an extremely low cytotoxic effect on standard cells. After being continuously cultured for 14 days, the surface of the Titanium scaffolds is fully covered with living cells which are stained by FDA, as seen by the green color in Fig. 6. Among these cells, only a minority of dead cells are observed. These dead cells are stained by PI, as seen by the red color in Fig. 6. This second cell culture study of seeding osteoblast (bone cells) onto the Titanium scaffolds has shown that the bone cells could attach on the scaffolds and that they could proliferate well within 14 days with minimum amount of dead cells. This shows that the Titanium implant prototypes are able to provide suitable substrate for cells to attach and live.
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In-vitro cell culture results have presented good and promising results. The fabricated porous Titanium dental implants are demonstrated to have very low toxicity effect. The implants have also shown that they can be used as suitable substrate for bone cells to live, grow and proliferate. In summary, as a bone-engineering material, the Titanium implant prototypes present good mechanical properties and biocompatibility by in-vitro investigation as stratum that supports cell adherence, cell proliferation and osteogenic differentiation, with little cytotoxicity effect. References
Fig. 5. Comparison of cytotoxicity between Titanium scaffolds and control cylinders.
Fig. 6. FDA and PI staining of cells on Titanium scaffolds 14 days after seeding.
4. Conclusions Commercially pure Titanium has been successfully investigated to fabricate porous dental implant prototype. Both the fabrication and cell culture results have been satisfactory. The dental implant prototype has been successfully fabricated using printing method and shown to have elastic modulus of 4.8–13.2 GPa. This elastic modulus is much lower than the modulus of the bulk commercially pure Titanium and is in the range of elastic modulus of natural bone [22]. The fabrication method also allowed variation of mechanical properties, as needed by the bone.
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