2019 Volume 60 Issue 11 Pages 2311-2318
A head-to-head comparison was made between an in-house-developed Ti–7.5Mo alloy and commonly-used grade-2 commercially pure titanium (CP-Ti) for dental casting applications. Experimental results indicated that all impurity concentrations, densities, linear thermal expansion coefficients and solid/liquid transition temperatures of Ti–7.5Mo and CP-Ti were similar. The 7-day total release of metal ions from immersed Ti–7.5Mo was acceptably low (10 µg/cm−2) with no harmful elements detected. Cytotoxicity testing indicated that Ti–7.5Mo had a cell viability of 82.5%, higher than the generally accepted value (70%). X-ray diffraction patterns indicated that Ti–7.5Mo was comprised primarily of α′′ phase with a small amount of β phase, while CP-Ti showed a monolithic α/α′ phase. Light and scanning electron microscopy revealed that CP-Ti had a typical plate-shaped morphology, while Ti–7.5Mo alloy was featured by its much finer acicular-shaped α′′ crystals along with equi-axed retained β grain boundaries. The castability value of Ti–7.5Mo was almost double that of CP-Ti. Both grinding and cutting tests indicated that Ti–7.5Mo had much better machinability than CP-Ti. Tensile testing indicated that Ti–7.5Mo had higher tensile strength, higher elongation and lower modulus (respectively 806 MPa, 42% and 70 GPa) than CP-Ti (respectively 571 MPa, 22% and 113 GPa). Bending data showed that Ti–7.5Mo had higher bending strength, lower modulus and much larger elastic recovery angle (respectively 1154.7 MPa, 75.8 GPa and 31.5°) than CP-Ti (respectively 919.5 MPa, 125 GPa and 2.8°). From all present data it was concluded that Ti–7.5Mo alloy would be a much better material than CP-Ti for removable partial denture application.
Fig. 4 Original wax pattern (a) and X-ray radiographs of RPD castings from CP-Ti (b) and Ti–7.5Mo alloy (c). A significant portion was missing (arrow) and internal voids (arrow) were detected in CP-Ti RPD.
Due to ever-increasing life expectancy, it is estimated that the portion of people over 65 will double its number to 10% of the world population by 2025.1) With the advance of oral hygiene, people are more susceptible to partial dentition rather than total denture.1,2) In the U.S., 43.7% of people over 20 have a tooth extracted and 43.1% of people over 65 have 6 or more teeth missing.2) Among the different forms of treatment, removable partial denture (RPD) remains to be an important option because it not only provides an effective solution for replacing missing teeth, but can also serve as provisional prostheses in case to overcome financial difficulties.2)
Traditionally, dental prostheses are often fabricated from alloys comprising high percentages of noble metals such as gold.2) Despite their good biocompatibility and corrosion resistance, the high costs of gold-based alloys have made them less attractive. Today’s RPD is most widely fabricated from Co–Cr based alloys. This metallic RPD has several advantages over acrylic resin RPD, such as higher thermal conductivity for a more natural feeling and higher mechanical strength which allows for design of thinner sections, thereby reducing the covering of the gingival margins. Nevertheless, there are still a number of drawbacks in Co–Cr RPD, such as high density (10 g/cm3), high rigidity and metallic taste. Keltjens et al.3) reported that most Co–Cr clasps became distorted and no longer fit the abutment correctly after 8 years of normal use. Co and Cr-related hypersensitivity issues have also been raised as early as the 1980s.4) Furthermore, the high modulus of elasticity (high rigidity) of Co–Cr alloy (about 210 GPa) is usually accompanied with smaller amounts of retentive undercut and larger amounts of healthy teeth being trimmed off, compared to the RPD made from a lower modulus material, such as titanium. This can be a disadvantage for Co–Cr RPD when esthetics or periodontal health is concerned.5) Co–Cr may also be connected to long-term bone loss reported by Isidor and Budtz-Jörgensen.6)
In view of all these disadvantages of Co–Cr alloy, commercially pure titanium (CP-Ti) for dental prosthesis applications has become attractive due to its low density (4.51 g/cm3), high corrosion resistance, high specific strength and high biocompatibility.7,8) The relatively low density of CP-Ti also allows porosities and other structural imperfections to be identified through conventional radiograph imaging.5) As mentioned above, the lower modulus of elasticity of CP-Ti allows larger amounts of retentive undercuts to be made than that recommended for Co–Cr alloy.5,9) Despite all these advantages over Co–Cr, the relatively low yield strength (YS) makes CP-Ti more susceptible to plastic deformation, especially in thin sections. Furthermore, its high melting point and high chemical reactivity make CP-Ti much harder to cast into conventional molds.10) The lower density of titanium also makes it hard to cast without a reasonable centrifugal force.11,12) To increase strength, Ti–6Al–4V alloy was also suggested for dental casting but the castability results were mixed, largely depending on the investment material used for casting.7,13) In addition, the release of Al and V ions from Ti–6Al–4V alloy has long been debated to possibly cause Alzheimer’s disease and cell cytotoxicity.14)
The authors’ laboratory has developed an Al/V-free, biocompatible, low modulus α′′-type Ti–7.5Mo alloy for orthopedic fixation applications. This alloy has a strength/modulus ratio significantly higher than the commonly-used Co–Cr alloy and CP-Ti. The excellent biocompatibility of Ti–7.5Mo alloy was demonstrated in an earlier study,15) wherein Ti–7.5Mo alloy with a modulus of elasticity of 78 GPa and Ti–6Al–4V alloy with a modulus of elasticity of 110 GPa were implanted into rabbit femur. The 26-week histology showed that the amount of new bone attached onto the Ti–7.5Mo implant was many times larger than that attached onto the Ti–6Al–4V implant. The better bone-implant interaction observed in Ti–7.5Mo was interpreted as a combined effect of chemistry (the absence of harmful Al and V) and modulus of elasticity (the much lower modulus of Ti–7.5Mo reduced the stress-shielding effect).
In the present study, a series of tests and measurements, including chemical composition, thermal expansion coefficient, solid/liquid transition temperature, ion release rate, cytotoxicity, castability and machinability, were conducted to evaluate the performance of the in-house-developed Ti–7.5Mo alloy for dental casting applications, particularly RPD. Furthermore, such properties as castability, machinability and mechanical properties, which are important factors for RPD performance, were compared between investment-cast Ti–7.5Mo alloy and the commonly-used grade-2 CP-Ti.
Both grade-2 CP-Ti (99.8 mass% pure) and Ti–7.5 mass%Mo alloy used for the study were purchased from China Steel Corporation (CSC), Kaohsiung, Taiwan. The concentrations of several most-concerned impurities, including O, C, N, H and Fe, are indicated in Table 1. The concentrations of N, C and H were provided by CSC, which were determined according to the methods set forth in ASTM E1409-13, ASTM E1941 and ASTM E1447, respectively. The concentration of O, provided by Société Générale de Surveillance (SGS), Taiwan, was measured according to ASTM E1409-13 method. The metal concentrations of Ti, Mo, Fe, Be, Cd and Ni were determined according to ASTM E2371 method using inductively coupled plasma–mass spectroscopy technique (ICP-MS-Element XR, Thermo Fisher, USA).
An ISO 22674 (“Dentistry-Metallic materials for fixed and removable restorations and appliances”) method was used to determine the metal ion release rates of Ti, Mo, Be, Cd and Ni. The corrosion-assisted ion release measurement was conducted by Light-Salt Testing Co. (Kaohsiung, Taiwan) using an inductively coupled plasma-optical emission spectrometry (ICP-OES) system (ICP-OES, Thermo Fisher Scientific Inc., USA) according to ISO 10271 method. The corrosion solution was prepared by adding 10 g of lactic acid and 5.85 g of sodium chloride into 300 ml of water, which was then diluted to 1000 ml having a pH value of 2.3. Samples for the test were cast into a 34 mm × 13 mm × 1.5 mm mold, ground and cleaned by ethanol and water in accordance to ASTM 600 method, followed by immersion in the corrosion solution for 7 days at 37°C. After the test, the corrosion solution was analyzed by ICP-OES for the ion concentrations of Be, Cd and Ni. The pH value of the corrosion solution was recorded and each average value was taken from 3 tests.
Cytotoxicity testing was conducted on Ti–7.5Mo alloy and CP-Ti using cultured NCTC clone 929 (BCRC 60091) cells. Samples for the test were immersed in a cell culture medium prepared by mixing 900 ml of α-MEM medium (Hyclone, MEM Alpha Modification with L-Glutamine, Ribo- and Deoxyribonucleosides), 100 ml of horse serum and 100 ml of three-in-one antibiotic (penicillin/streptomycin/fungizone) at a ratio of 1 g sample/5 ml (culture medium) for 24 h at 37°C. The negative control group was prepared by immersing Al2O3 balls in the same cell culture medium with the same ratio (1 g/5 ml), temperature and time. The positive control groups were prepared from three different concentrations (0.01, 0.1 and 0.3%) of phenol dilute, which were prepared by respectively adding 1.0, 10 and 30 µl of phenol mixed with 10 ml of cell culture medium under the same condition. Cells were plated at 5 × 103 cells/well with cell culture medium. After being incubated in 5% CO2 air at 37°C for 24 h, the old cell culture medium was replaced by the extract of Ti–7.5Mo alloy, blank medium and both negative and positive control groups for 24 h at 37°C. After the second incubation, the extract solution, blank medium and both negative and positive control groups were removed from each well and the well was fed with 100 µl new cell culture medium/well. 10 µl of water-soluble tetrazolium salt (WST-1) was added to each well and the cells were incubated for 1 h. The viability of cells was determined by WST assay with an ELISA reader at an absorbance of 450 nm.
The density of Ti–7.5Mo alloy was determined by directly measuring the volume and weight of Ti–7.5Mo ingots. The linear thermal expansion coefficient of Ti–7.5Mo alloy was determined by differential thermal analysis (DTA) (Labtronic II, Theta Industries Inc., USA) according to ISO 22674 method. The sample for the test was cast into a 3 mm dia., 20 mm high cylindrical mold. The temperature range for the test was between room temperature and 550°C at a heating rate of 5°C/min. The linear thermal expansion coefficient of CP-Ti was reported to be 8.4–10.8 × 10−6/°C.16)
The liquid/solid transition temperature of Ti–7.5Mo alloy was measured by Fosan Nanhai Jingmei Testing Service Co. according to the method set forth in ISO 22674. Samples for the test were cast into a triangular cone-shaped mold with a width of 8.5 mm and height of 30 mm, wherein the cone had an intersection angle of 8° with the vertical line. The test sample was heated to 200°C below the theoretical transition temperature (1740°C) at a heating rate of 2.5°C/min until the cone bent and the tip of the cone touched the platform. Such temperature was considered as the liquid/solid transition temperature of the sample.
Investment casting was performed using lost wax casting technique using a centrifugal casting machine (Ticast Super R, Selec, Japan), wherein an aluminum oxide-based investment material (TiM, Selec, Japan) was used for the study. A mesh-shaped wax pattern was used to evaluate and compare the castability of CP-Ti and Ti–7.5Mo alloy. The metal ingot was melted in an open-based graphite hearth in argon gas, before pouring into the mold assisted by the centrifugal force of the casting system. After casting, the mold material was removed by sand blasting and the cast sample was polished to meet each test requirement. To further compare the quality of CP-Ti and Ti–7.5Mo castings, two RPDs with the same wax pattern design, same material weight and same casting parameters were prepared and radiographically analyzed. A non-destructive XT V160 X-ray system (Nikon Co., Japan) was used at 125 kV and 75 µA for 500 ms.
To assess machinability of Ti–7.5Mo alloy and CP-Ti, which is a practically important factor to dentists as well as dental labs, two different types of tests, grinding and cutting, were performed. The cutting times of Ti–7.5Mo alloy and CP-Ti with the same cross-section area (1.3 mm × 5 mm) were determined using a low speed diamond saw (ISOMet®, Buehler Inc., USA) under a load of 100 g and a speed of 300 RPM (equivalent to a linear speed of 120 m/min). The average cutting time was calculated from 6 repeated tests under the same condition. The grinding test was conducted using SiC sandpapers. The initial surface roughness of Ti–7.5Mo alloy and CP-Ti was maintained the same by polishing both materials using #100 sandpaper. The #100 sandpaper-polished 10 mm × 10 mm plate samples were further polished for 60, 120 and 180 seconds on #1200 sandpaper. The grindability was evaluated by the time taken to reach the same final surface roughness. The average surface roughness of the sample was taken from six repeated tests under the same condition.
X-ray diffraction (XRD) for phase analysis was conducted using a Bruker D2 Phaser diffractometer (Bruker Co., Germany) operated at 30 kV and 10 mA with scanning speeds of 2°/min and 0.1°/min. A Ni-filtered CuKα radiation was used for the study, while a silicon standard was used for the calibration of diffraction angles. The various phases were identified by matching each characteristic peak in the diffraction patterns with JCPDS files (Joint Committee on Powder Diffraction Standards, now called International Centre for Diffraction Data, ICDD).
Microstructural examination of the cast samples was performed using an optical microscope (Leica TMX 100, Germany). The surfaces of the materials for light microscopy were mechanically polished via a standard metallographic procedure to a final level of 0.05 µm alumina powder, followed by chemical etching in a mixture of water, nitric acid, and hydrofluoric acid (100:3:1 by volume). A scanning electron microscope (SEM) (JEOL JSM-6510, Japan) operated at 5 kV under secondary electron mode was also used for microstructural examination in more detail. The samples for SEM examination were prepared under the same procedure as for optical microscopy.
Vicker’s hardness of Ti–7.5Mo alloy was measured according to ASTM E-92 method using a microhardness tester (HMV-G20S, Shimadzu Scientific Instruments Co., Japan) with a loading force of 1 kgf and residence time of 15 sec. Samples for the microhardness test were prepared under the same procedure as for optical microscopy. The average microhardness values were taken from six samples at two different locations.
To evaluate and compare the mechanical properties of the two materials, 3-point bending and tensile tests were performed. The 3-point bending test was conducted using a Shimadzu AGS-500D mechanical tester (Japan). The bending strength was determined from the equation, σ = 3PL2bh−2, were σ is the bending strength (MPa), P is load (kg), L is span length (mm), b is the specimen width (mm) and h is the specimen thickness (mm). The dimensions of the specimens for testing were: L = 30 mm, b = 5 mm and h = 1 mm. The modulus of elasticity in bending were calculated from the load increment and the corresponding deflection increment between the two points on the straight line as far apart as possible using the equation, E = L3δP4bh−3δc, where E is the modulus of elasticity in bending (Pa), P is the load increment as measured from the pre-load (N) and δc is deflection increment at mid-span as measured from pre-load. The average bending strength and modulus of elasticity were taken from six tests. The elastic recovery (“bounceback”) capability was evaluated from the change in deflection angle when the load was removed. The tensile test was conducted using a servo-hydraulic type testing machine (AG-10KNX, Shimadzu, Japan) according to ASTM E8 method. The testing was performed at room temperature with a constant crosshead speed of 1.5 mm/min. The average ultimate tensile strength (UTS), yield strength (YS) at 0.2% offset, modulus of elasticity, and elongation to failure were taken from six tests under each condition.
The chemical compositions of Ti–7.5Mo alloy and grade-2 CP-Ti used for the study are listed in Table 1. The average Mo content of Ti–7.5Mo alloy was measured to be 7.48 mass%, quite close to the preset goal (7.5 mass%). The concentrations of the most-concerned impurities in titanium/titanium alloys, O, Fe, C, N and H, were measured to be 0.15, 0.034, 0.015, 0.012 and 0.0004 mass%, respectively. Compared to the commercial binary beta-phase Ti–15Mo alloy (ASTM F2066-Standard Specification for Wrought Titanium-15 Molybdenum Alloy for Surgical Implant Application), all the impurity levels of the present Ti–7.5Mo alloy were within the ranges required for Ti–15Mo alloy. Table 1 also indicates that the concentrations of C, O, N and H of Ti–7.5Mo were within the ranges required for grade-2 CP-Ti set forth in ASTM B265-15 (Standards Specification for Titanium and Titanium Alloy Strip, Sheet and Plate).
3.2 Density, thermal expansion coefficient and solid/liquid transition temperatureSome physical properties closely related to the casting performance of metals, including density, thermal expansion coefficient and solid/liquid transition temperature, were determined for Ti–7.5Mo alloy in this study. The density of the orthorhombic Ti–7.5Mo alloy, 4.7 g/cm3, was slightly higher than pure Ti (4.5 g/cm3) due to the presence of the heavier Mo atoms.
Thermal expansion coefficient is a crucial factor in lost wax casting procedure because it is closely related to the dimensional accuracy of cast articles. The thermal expansion coefficient of the metal should be as close as that of the investment as possible. The measured linear thermal expansion coefficient of the present Ti–7.5Mo alloy was 9.7 × 10−6 K−1, which was close to the linear thermal expansion coefficient of CP-Ti (9.6 × 10−6 K−1 at 25–500°C) and Ti–6Al–4V (10.2 × 10−6 K−1 at 25–500°C).
According to the binary Ti–Mo phase diagram and theoretical melting points of pure Ti (1668°C) and pure Mo (2623°C), the solid/liquid transition temperature of Ti–7.5Mo alloy should be near 1740°C. The presently measured solid/liquid transition temperature of Ti–7.5Mo alloy turned out to be about 1700°C, quite close to the calculated solid/liquid transition temperature for this high temperature measurement. The little higher solid/liquid transition temperature of Ti–7.5Mo alloy (1700°C) than pure Ti (1668°C) would cause a little lower overheat during melting, which in principle could cause the viscosity of the molten metal to somewhat decrease and castability to also decrease. Nevertheless, as will be shown later, the present Ti–7.5Mo alloy demonstrated surprisingly higher castability than CP-Ti, indicating that there must be other factors contributing the much better castability observed in Ti–7.5Mo alloy than pure Ti, as will be further discussed in a later section.
3.3 Ion release rateThe release of metal ions from RPD is crucial from clinical point of view. According to the aforementioned ISO 22674 method, after being immersed in the specified corrosion solution for 7 days, the amount of total metal ions released from the test sample into the corrosion solution should not be over 200 µg/cm2 and no harmful elements should be detected. The present test showed that, after Ti–7.5Mo sample was immersed for 7 days, the total release of metal ions as only 10 µg/cm2 with no harmful elements detected.
ISO 22674 further requires that the amounts of Be and Cd should be each less than 0.2 mass%, while Ni should be labeled if it has an amount higher than 0.1 mass%. As described in Methods, both direct chemical analysis (ICP-MS) and corrosion-assisted ion release measurement (ICP-OES) were conducted in this study. The results indicated that, in each test, all Be, Cd and Ni concentrations were below the detection limit, i.e., <1 µg/cm2, demonstrating that the present Ti–7.5Mo alloy would not release any of such harmful elements. It is also worth noting that Ti–7.5Mo alloy does not contain low biocompatibility metallic elements such as Co, Cr and V, which are present in Ti–6Al–4V or Co–Cr alloy.
3.4 Cytotoxicity testThe result of cytotoxicity test revealed that the present Ti–7.5Mo alloy had a cell viability of 82.5%, which was higher than the generally accepted value (70%). According to ISO 10993-5, it can be concluded that Ti–7.5Mo alloy would not have a negative effect to human cells.
3.5 Crystal structure and microstructureThe high speed (2°/min) and low speed (0.1°/min) XRD patterns of cast Ti–7.5Mo alloy and CP-Ti are shown in Figs. 1(a) and (b), respectively. The XRD patterns clearly indicated that the cast Ti–7.5Mo alloy was comprised primarily of α′′ phase with an orthorhombic crystal structure, along with a small amount of β phase with a body-centered cubic (bcc) crystal structure, consistent with earlier reports from the present authors’ laboratory. The cast CP-Ti, on the other hand, had monolithic α/α′ phase. The term “α/α′ phase” or “α/α′ peaks” was used in this study due to the fact that α phase and α′ phase have the same hexagonal close-packed (hcp) crystal structure and are indistinguishable from XRD patterns. These two phases are often distinguished from each other by their different morphologies. Typically α′ phase has a fine, martensitic-type acicular morphology usually obtained from a fast cooling process, while α phase usually exhibits a coarser plate-shaped morphology.17)
XRD patterns of Ti–7.5Mo alloy and CP-Ti (a) scan speed 2°/min; (b) scan speed 0.1°/min.
As shown in Figs. 2(a) and (b), the cast CP-Ti had a typical plate-shaped morphology (with a plate width of about 1.0–2.0 µm). The cast Ti–7.5Mo alloy was primarily featured by its much finer acicular-shaped α′′ crystals (with a width of about 0.1–0.3 µm), along with essentially equi-axed grain boundaries which were retained as the alloy went through β phase field during the rapid cooling (casting) process. Nevertheless, the XRD-revealed retained β phase was not identifiable from these optical and scanning electron micrographs.
Optical (a), (c) and SEM (b), (d) micrographs of cast CP-Ti (a), (b) and Ti–7.5Mo alloy (c), (d).
To compare the castability of Ti–7.5Mo alloy and CP-Ti, a mesh type wax pattern was used for the study. Mesh pattern has been popularly used to assess metal/alloy castability for dental casting applications, especially RPD.13,14) The rectangular-shaped wax used was 0.3 mm in thickness and 3525 mm2 in area. The average castability value was obtained from six repeated tests under the same condition. For comparison, the castability of grade-2 CP-Ti was also measured using the same mold design under the same casting condition. The result indicated that the mold-filling ratio of Ti–7.5Mo alloy was almost double that of CP-Ti (Fig. 3). The superior quality of Ti–7.5Mo RPD casting was also demonstrated in Fig. 4. As shown in the X-ray radiographs, not only a significant portion was missing, internal voids were detected in CP-Ti RPD. On the other hand, the casting of Ti–7.5Mo RPD was complete without radiographically detectable voids.
Mold-filling ratios (a) and typical cast samples (b) of Ti–7.5Mo alloy and CP-Ti.
Original wax pattern (a) and X-ray radiographs of RPD castings from CP-Ti (b) and Ti–7.5Mo alloy (c). A significant portion was missing (arrow) and internal voids (arrow) were detected in CP-Ti RPD.
In general, adding alloy elements into a pure metal often reduces the castability of the metal.13,14) However, it was found in this study that the addition of 7.5 mass% Mo into Ti largely enhanced castability. A number of factors could affect the castability of a metal, such as casting environment, mold design, mold temperature, cooling rate, molten metal temperature, viscosity and its reaction with mold material, surface tension (a measure of surface free energy), dendrite formation, among others.14,18) Since all the casting process parameters remained the same in this study, such extrinsic factors affecting castability as mold design, mold temperature, etc. could be ruled out. The factors most likely to affect castability in this study would be material-related factors, such as mold reaction, viscosity, surface tension and dendrite formation. It was reported that addition of alloy elements to a pure metal generally decreased the fluidity and castability of the metal due to dendrite formation disrupting the molten metal flow.18) Apparently this dendrite factor could not either explain the much better castability of Ti–7.5Mo over CP-Ti. Although the centrifugal force during casting favors Ti–7.5Mo alloy which has a higher density than CP-Ti due to the presence of Mo, whether this slight increase in density (by only 4.3%) could cause the drastic increase in castabilty is hard to fathom and needs further clarification.
3.7 MachinabilityAs described in Methods, two simple tests (grinding and cutting) were performed to assess machinability of Ti–7.5Mo alloy and CP-Ti. The cutting test result indicated that the average cutting times per unit area for Ti–7.5Mo alloy and CP-Ti were 69.6 and 80.5 sec/mm2, respectively. The grinding test result indicated that the average grinding times from an initial surface roughness of 0.6 µm to a final roughness of 0.3 µm for Ti–7.5Mo alloy and CP-Ti were about 60 sec and 180 sec, respectively. The much better machinability of Ti–7.5Mo alloy than CP-Ti may be a direct result of the fine morphology and higher hardness of Ti–7.5Mo alloy.
3.8 Hardness, bending and tensile propertiesThe average Vickers hardness value of cast Ti–7.5Mo alloy was measured to be 369.3 ± 21.4 HV, while the average Vickers hardness value of grade-2 CP-Ti was 186.4 ± 6.3 HV. Revised ANSI/ADA Specification No. 5 for Dental Casting Alloys requires that the standard deviation of hardness should be less than 10%. The standard deviation of the present cast Ti–7.5Mo alloy was 5.8%, within the ADA-required standard deviation range. The observed higher hardness of Ti–7.5Mo alloy might be a direct result of Mo-induced solution hardening effect.6) The cross-sectional microhardness measurement indicated that the α-case thicknesses for both CP-Ti and Ti–7.5Mo alloy were roughly 50–100 µm.
The bending properties of cast Ti–7.5Mo alloy and grade-2 CP-Ti are shown in Fig. 5. As indicated in the figure, the bending strength of Ti–7.5Mo alloy (1154.7 MPa) was much higher than that of CP-Ti (919.5 MPa), while the bending modulus of Ti–7.5Mo alloy (75.8 GPa) was much lower than that of CP-Ti (125.0 GPa). The lower bending modulus of Ti–7.5Mo alloy was a direct result of the presence of orthorhombic α′′ phase which has an inherently low modulus.18) The combination of the higher strength and lower modulus gave rise to a drastically larger elastic recovery angle of Ti–7.5Mo alloy than CP-Ti (31.5° vs. 2.8°). The large elastic recovery angle of Ti–7.5Mo alloy can make the material an ideal candidate for clasp.
Bending properties of cast Ti–7.5Mo alloy and CP-Ti. (a) Bending strength and modulus; (b) typical load-displacement profiles; (c) illustration of bounceback angle measurement.
Figure 6 compares the tensile properties between cast Ti–7.5Mo alloy and cast CP-Ti. Both materials had a similar YS (405 MPa), while the ultimate tensile strength (UTS) of Ti–7.5Mo alloy (806 MPa) was much higher than that of CP-Ti (571 MPa). Cast Ti–7.5Mo alloy also demonstrated a much larger elongation (42%) than cast CP-Ti (22%). Similar to that observed in bending test, the tensile modulus of Ti–7.5Mo alloy (70 GPa) was far lower than that of CP-Ti (113 GPa). Again, the much lower tensile modulus of Ti–7.5Mo alloy was a direct result of the presence of low-modulus α′′ phase. The higher tensile strength of Ti–7.5Mo alloy could be a combined result of solution strengthening effect and its finer morphology of the alloy.
Tensile properties of cast Ti–7.5Mo alloy and grade-2 CP-Ti. (a) Tensile strength and elongation; (b) tensile modulus.
