Phases percentages and total porosity of TNS and TNS/M samples sintered at 900°C and 1100°C.
\r\n\tHomeostasis is brought about by a natural resistance to change when already in the optimal conditions, and equilibrium is maintained by many regulatory mechanisms. All homeostatic control mechanisms have at least three interdependent components for the variable to be regulated: a receptor, a control center, and an effector. The receptor is the sensing component that monitors and responds to changes in the environment, either external or internal. Receptors include thermoreceptors and mechanoreceptors. Control centers include the respiratory center and the renin-angiotensin system. An effector is a target acted on to bring about the change back to the normal state. At the cellular level, receptors include nuclear receptors that bring about changes in gene expression through up-regulation or down-regulation and act in negative feedback mechanisms. An example of this is in the control of bile acids in the liver.
\r\n\tSome centers, such as the renin-angiotensin system, control more than one variable. When the receptor senses a stimulus, it reacts by sending action potentials to a control center. The control center sets the maintenance range—the acceptable upper and lower limits—for the particular variable, such as temperature. The control center responds to the signal by determining an appropriate response and sending signals to an effector, which can be one or more muscles, an organ, or a gland. When the signal is received and acted on, negative feedback is provided to the receptor that stops the need for further signaling.
\r\n\tThe cannabinoid receptor type 1 (CB1), located at the presynaptic neuron, is a receptor that can stop stressful neurotransmitter release to the postsynaptic neuron; it is activated by endocannabinoids (ECs) such as anandamide (N-arachidonoylethanolamide; AEA) and 2-arachidonoylglycerol (2-AG) via a retrograde signaling process in which these compounds are synthesized by and released from postsynaptic neurons, and travel back to the presynaptic terminal to bind to the CB1 receptor for modulation of neurotransmitter release to obtain homeostasis.
\r\n\tThe polyunsaturated fatty acids (PUFAs) are lipid derivatives of omega-3 (docosahexaenoic acid, DHA, and eicosapentaenoic acid, EPA) or of omega-6 (arachidonic acid, ARA) and are synthesized from membrane phospholipids and used as a precursor for endocannabinoids (ECs) mediate significant effects in the fine-tuning adjustment of body homeostasis.
\r\n\t
\r\n\tThe aim of this book is to discuss further various aspects of homeostasis, information that we hope to be useful to scientists, clinicians, and the wider public alike.
Recrystallization of metal material is a complicated phenomenon that unites different aspects of its profile, including plastic deformation, strain hardening and behaviour by heat treatment. The latter entails numerous processes such as recovery, growth of recrystallization nuclei and the competition of new grains. When considering the plastic deformation of single crystals, the experimental analysis of their ensuing recrystallization becomes simpler: in the grains of a polycrystal, plastic deformation creates a wide spectrum of substructures depending on the crystallographic orientation of grains, whereas in the case of a deformed single crystal, initial substructure conditions are much more limited.
In principle, concepts of \'polycrystal\' and \'single crystal\' ought to be associated with the concept of that of \'deformed single crystal\'. The single crystal, subjected to plastic deformation (most often one deals with rolled single crystals) does not become a polycrystal; at the same time, it ceases to be a single crystal, because of the crystallographic texture that forms in it due to the number of initial orientations of this single crystal. A rolled polycrystal is a conglomerate of rolled single crystals, but a rolled single crystal differs from a polycrystal (or a rolled polycrystal) in terms of boundary characteristics, which separate regions with different orientations. In the polycrystal, these regions correspond to individual grains and are separated from neighbours by high-angular boundaries. In the rolled single crystal, these regions originate from the same monocrystalline matrix, due to the inevitable non-uniformity of plastic deformation; therefore, they are separated from neighbours by low-angular boundaries, except for the case of twinning, which was most often observed in the current work.
Experiments involving the rolling of single crystals can provide results that may be unreachable by similar experiments with polycrystals. Among such results is the obtaining of quasi-monocrystalline foils by the cold rolling of BCC single crystals (Mo, Nb) in the case of initial orientation {001}<011>. These foils are characterized by minimal strain hardening up to the highest deformation degrees and do not recrystallize by subsequent annealing [1]. At the same time, rolled single crystals with other initial orientations acquire significant strain hardening and can even be destroyed under rolling. These effects make clear the mechanisms of plastic deformation and recrystallization of materials, and indicate how the substructure of polycrystalline products can be modified and optimized. To date, the existence of analogous effects in the rolling of α-Zr single crystals have not been investigated due to difficulties related to procuring them.
Though the title of this chapter is “Recrystallization of rolled α-Zr single crystals”, the intention of authors is wider and concerns also recrystallization of rolled polycrystalline α-Zr, since the behaviour under annealing of rolled single crystals helps to explain some known features of recrystallization in polycrystals and can compare the characteristics of strain hardening distribution in rolled single crystals and polycrystals.
When considering recrystallization and especially its textures, it is necessary to take into account the prehistory of the material, including features of its previous deformation, grain reorientations, formed deformation textures and plastic deformation mechanisms. Therefore analysis of recrystallization inevitably expands in scope and covers a number of adjacent questions, among them, the details concerning preceding grain reorientation. Thus, in the current chapter, recrystallization is considered as a derivative effect of plastic deformation.
As a preamble to discussing the behaviour of α-Zr single crystals under rolling, we take into account the Hobson’s calculated diagram for operating plastic deformation mechanisms [2] and our X-ray experimental observations of basal axes (or axis) trajectories during the course of rolling [3, 4]. The Hobson’s diagram, added later by Matcegorin et al. [5] (Fig. 1), predicts slip and twinning systems operating in α-Zr grains in conditions of cold rolling. Three twinning systems are considered: (1) {10
Additionally, it was found [6] that development of the rolling texture in polycrystalline α-Zr has three stages:
formation of the quasi-stable texture T1 (0001)±15-20°ND-RD <11
transition T1
perfection of the stable texture T2 (0001)±30-40°ND-TD <10
Texture T1 is supported due to an equilibrated slip at the basal and pyramidal planes, transition T1
Regions of stereographic projection for the α-Zr sample, whereby its rolling, different deformation systems operate depending on the orientation of the grain basal axis. Calculated systems are characterized by the largest values of the Schmidt factor. Assumptions were made that each system has the same critical stress and that σRD = σTD[
Development of the rolling texture in α-Zr single crystals was studied as applied to coarse-grained samples [7], since lattice reorientation in separate coarse grains as a first approximation occurs independently of similar processes in other grains. However, the formation of the recrystallization texture in the same coarse-grained sample includes the growth of nuclei at grain boundaries, resulting in this process moving outside of separate grains and its control via the orientations of neighbouring grains. Therefore, recrystallization in rolled coarse-grained samples does not imitate the situation in rolled single crystals. Thus, the current study is the first X-ray investigation of recrystallization in cold-rolled α-Zr single crystals.
As for recrystallization of cold-rolled polycrystalline α-Zr, its regularities were described in detail in [8-9]: reorientation of prismatic axes by 30° about the basal axis, the shift of basal axes along the TD-ND diameter of stereographic projection in the direction of ND and in the case of multicomponent rolling, texture redistribution of its components according to definite relationships. However, these data were obtained only by the geometric analysis of texture features, without the use of data concerning the distribution of strain hardening in rolled polycrystals. Therefore, the motive power of recrystallization in the considered concrete cases is unclear. Currently, the X-ray method of generalized pole figures is being developed [10, 11], which allows revealing the distribution of strain hardening in textured materials, depending on the crystallographic orientation of their grains. This will allow a better understanding of the reasons for the observed passage of recrystallization.
Examples of the experimental trajectories of basal axes reorientations during the course of cold rolling of Zr and its alloys.
Evolution of PF (0001) for α-Zr by increasing the deformation degree through cold rolling.
Calculated trajectories of basal axes reorientations during the course of rolling through the operation of basal (a) and pyramidal (b) slip systems [
A single crystal of pure Zr was obtained by means of solid-phase over-crystallization. Samples 7x10x3 mm in size were cut from the cylindrical single crystal using the electro-erosion method. Obtained samples were rolled at a laboratory rolling mill between plates of stainless steel up to the deformation degree εmax= 80% and with ~5% reductions per pass.
The initial monocrystalline plate and rolled single crystals were subjected to X-ray texture investigation, which included (1) texture analysis using the method of direct pole figures; (2) the study of the strain hardening distribution using the method of generalized pole figures (GPF) (see the next section).
Then, rolled single crystals were annealed in the evacuated vessel at 580oC for 3 h, the standard regime of recrystallization heat treatment for industrial products made from Zr-based alloys. After annealing, rolled single crystals were repeatedly subjected to X-ray investigation; direct pole figures were measured once more and subtraction diagrams were constructed on the basis of texture data treatment (see the next section).
In addition to the investigation of rolled α-Zr single crystals, the current work also included the study of cold-rolled and recrystallized (1) coarse-grained iodide Zr and (2) a polycrystalline sheet of Zr-1%Nb alloy, rolled along RD and along TD. The development of rolling textures, distributions of strain hardening and recrystallization textures in these samples was studied using the same X-ray methods as those used in the study of rolled single crystals. When comparing X-ray data obtained for rolled single crystals and polycrystalline samples, a number of important conclusions concerning the regularities of their strain hardening and recrystallization were made.
All X-ray studies of the used samples were carried out using an X-ray diffractometer Bruker D8 DISCOVER with a LynxEye position-sensitive detector. Partial direct pole figures PF(0001) and {11
For rolled α-Zr single crystals, the X-ray method of generalized pole figures (GPF) was used. This method consists of registration of X-ray line profiles during the course of texture measurement [10, 11]. In the current case, the method characterized the condition of the crystalline lattice along basal axes <0001> in all grains of the sample. The physical angular half-width β0002 of the X-ray line (0002), increasing as a result of strain hardening and decreasing by recrystallization was used as a measure of crystalline lattice distortion. A wide spectrum of substructure conditions, i.e., values of β0002, the distribution of which in the stereographic projection is shown with GPF β0002, corresponded to separate micro-fragments of the rolled single crystal or grains of rolled polycrystal, and the optimal criterion for systematization of GPF data was their crystallographic orientation.
In particular, for many rolled metal materials with the developed texture it was revealed that the strain hardening of fragments (grains, subgrains, crystallites) corresponding to texture maxima was significantly lower than that for texture minima. Strain hardening increased as grain orientation shifted from a texture maximum to a texture minimum. In cold-rolled BCC metals, where primary recrystallization is realized by the growth of grains with increased strain hardening, new maxima of the recrystallization texture grew at the maxima “slopes” of the rolling texture. Hence, strain hardening of these grains was sufficiently high for their predominant growth; at the same time, there were a sufficient amount of grains with such orientation for them to be able to “swallow” the rest of the deformed matrix.
Application of the GPF method requires the availability of an X-ray texture diffractometer of last models with a position sensitive detector, which will allow registering the entire X-ray line profile by texture measurement, without time losses for movement of the detector. The procedure of GPF measurement and construction requires special software, which can be developed on the basis of the standard technical potential of BRUKER diffractometers.
In order to reveal regions of pole density changes in PF(0001) through the transition from cold-rolled to recrystallized α-Zr, difference diagrams S(ψ,φ) are constructed in which (ψ,φ) are the angular coordinates of points in PF and S(ψ, φ) is equal to the difference between pole densities in corresponding points of PF(0001) for recrystallized and cold-rolled samples; thus, S(ψ, φ) = [PF(0001)recr-PF(0001)roll]. Inhomogeneous strain hardening and recovery result in some violation of proportionality between scattering ability, calculated pole density and specific volumes of deformed and recrystallized grains. In this connection, pole densities in the PF of rolled and recrystallized samples can be compared only with limited accuracy. Therefore, subtraction diagrams are constructed only for those regions of PF(0001)roll where pole densities noticeably exceed the sensitivity of texture measurements. In the current case, this threshold value of pole density was accepted as 0.5. Represented below, subtraction diagrams visually demonstrate the arrangement of regions with maximal changes of pole density as a result of recrystallization.
In Fig. 5a, the partial direct PF{11
The same cycle of investigations was repeated for monocrystalline plates of three initial orientations, differing only in terms of RD. These orientations corresponded to the following angular coordinates of basal axes: N 1 - (70°, 33°); N 2 - (70°, 152°); N 3 - (70°, 87°). The three columns of PFs in Fig. 5 demonstrate the successive stages of rolling texture development in single crystals within the following indicated orientations: N 1 - (
After deformation by only 5-7%, the texture maximum in PF(0001) characterizing the initial orientation of a single crystal broke into five-six new maxima corresponding to the reorientation of single crystal fragments by means of twinning in planes {10
The above-described changes of basal axes orientations when rolling α-Zr single crystals with orientations N 1 and N 2 agreed with the proposed diagram by Hobson of plastic deformation mechanisms activated in α-Zr grains with different orientations of basal axes relative to external directions ND, RD and TD [2, 5]. The process of twinning continues until where some regions with the initial orientation remain in the single crystal.
This process is described using curves that show changes in PF parameters as the deformation degree increases (Fig. 6). Among these parameters, there are angular positions ψ and intensities P of initial texture maxima
At the next stages of deformation in the single crystal N 2, twinning by planes {11
Thus, up to ε=30% texture formation under rolling was provided by the active operation of twinning by planes {10
PF{11
Fig. 5b and c, and Fig. 6c show that twinned areas of single crystals (fragments separated from other parts of a single crystal by planes of twinning) shift gradually to the centre of PF(0001), i.e., to ND of the rolled plate. Such a shift of texture maxima is conditioned by the development of a basal slip in the material. Along with clearly expressed additional texture maxima arising (see Fig. 5j), situated at the diameter ND-TD and symmetrically relative to ND, and connected with twinning by planes {11
Changes in the pole coordinate ψ (a, c) and intensity (b, d) of texture maxima
Results of the X-ray study of recrystallization in rolled α-Zr single crystals are presented in Fig. 7 and 8 for single crystal N 1, in Fig. 9 and 10 for single crystal N 2 and in Fig. 11 and 12 for single crystal N 3. The texture PF(0001) for rolled samples and recrystallized samples are shown in these figures as (
The X-ray method of GPF, used for rolled single crystals in the current case characterized the condition of the crystalline lattice along basal axes <0001> in all fragments of the sample. As a measure of lattice distortion, which strengthens as a result of strain hardening and is removed by recrystallization, the physical angular half-width of the X-ray line (0002) β0002 was applied. A wide spectrum of substructure conditions, i.e., values of β0002, responded to separate fragments of the rolled single crystal. In subtraction diagrams S(ψ, φ) = [PF(0001)recr- PF(0001)roll] for each point (ψ, φ) of the stereographic projection, the difference between the pole densities of recrystallized and a rolled single crystal was calculated. The presented subtraction diagrams visually demonstrated the arrangements of regions where maximal changes of pole density by recrystallization were observed. Gradations of red colour in the subtraction diagram indicate regions of positive values and respond to an increase in pole density as a result of recrystallization, whereas gradations in blue colour correspond to a decrease of pole density due to recrystallization. Diagrams of correlation between PF(0001) and GPF β0002, shown in Figs. 7e and 8e, had an appearance typical to that of a first approximation for many deformed metal materials, characterized by a gradual decrease of values β0002 as the peak of the texture maximum approached. Against a background of this tendency features can be observed as clearly connected to the character of corresponding rolling textures.
Due to the use of rolled α-Zr single crystals, the rolling textures obtained were formed mainly by twinning. In particular, in a single crystal N 1 rolled at 53%, i.e., by means of twinning, the rolling texture arose with maxima, which in spite of a relatively high deformation degree, took up transition positions, distinguishing them from maxima in stable textures T1 and T2. Subtraction diagrams showed that in the case of recrystallization within regions of PF for a rolled single crystal, where texture maxima arose due to twinning, the growth of pole density was absent (Fig. 7a, d). In fragments of rolled single crystals corresponding to the orientation of their basal axes near normal direction ND, rolling can activate only a pyramidal slip, characterized by high values of critical shear stress. Operation of other slip systems in these fragments (or grains of polycrystal) is suppressed due to low Schmidt factor values. Therefore, in fragments of α-Zr single crystals (or the grains of a polycrystal) experiencing reorientation, predominantly through the participation of twinning, their strain hardening, including distortion of the crystalline lattice and fragmentation of substructure elements, is relatively small. So, during the course of subsequent annealing, nuclei of recrystallization do not arise in these regions and these fragments (or grains of polycrystal) are swallowed up by recrystallization nuclei, growing more actively and corresponding by their orientation to regions of PF with increased strain hardening. GPF β0002 for rolled single crystals N 1 in Figs. 7c and 8c also indicate that the strain hardening of fragments with a basal axis orientation near ND is minimal. For this reason, in subtraction diagrams constructed for all rolled single crystals, we see a blue centre, testifying that in the central region the growth of pole density by recrystallization was absent, since the only mechanism provided here in terms of its input was twinning.
Recrystallization in cold-rolled α-Zr single crystal N 1 (ε~50%):
Recrystallization in cold-rolled α-Zr single crystal N 1 (ε~80%):
At the same time, the subtraction diagrams constructed for rolled single crystals show that its texture maxima are bordered by zones of intensive recrystallization (Fig. 7d). Only within these regions of stereographic projection was the increased lattice distortion also observed (Fig. 7c) as a consequence of the above-described simultaneous operation of the pyramidal and basal slip. According to GPF β0002, the “slopes” of texture maxima (see Fig. 7c) respond to regions of greatest fragmentation and the distortion of the crystalline structure, so that in the current case, the main mechanism of α-Zr recrystallization was the intensive growth of regions in the deformed crystalline matrix with the increased energy of residual distortions.
As the deformation degree of single crystals increased, their rolling texture changed and acquired features of the texture T2 (Fig. 7-12). This testified that the relative input of twinning in the formation of new maxima gradually decreased and was replaced by the slip. When this happened, the character of subtraction diagrams and diagrams of correlation PF(0001) – GPF β0002 sharply changed. In the subtraction diagrams, additional red regions arose, coinciding approximately with the positions of T2 maxima in rolling textures. Hence, the strain hardening at these maxima was higher than in regions where twinning had previously been predominant and here, a greater number of recrystallization nuclei arose (Fig. 8, 10, 12).
The most detailed and precise data concerning the evolution of strain hardening distributions in rolled single crystals provided diagrams of correlation PF(0001) – GPF β0002, showing in Figs. 7e and 8e. Texture maxima, formed by twinning, were twice as high as those formed by the slip at a later stage. Strain hardening of fragments corresponding to these maxima by ε=53% was constant and remained low (β0002< 0,2o) for the most part, whereas for ε=80% strain hardening as noticeably higher (β0002> 0,25o) and varied within significant intervals. The correlation diagrams show that strain hardening grew as an angular distance of the considered fragment from the top of maxima increases. The reason for this regularity was the different features of the slip within the stable texture maxima and at their periphery parts, i.e. at texture minima, where pole density is lower than 1. Stable texture maxima arose in regions of stereographic projection, where the operation of different slip systems was equilibrated. In the case of α-Zr rolling stability of T2, texture maxima were conditioned through the mutually balanced operation of basal and prismatic slip systems. However, the joint activity of these slip systems resulted in increased strain hardening, whereas their alternating operation, realized in turn, that is one after another, led to a drop in strain hardening and was equivalent to the effect of the monoslip. The metal matrix adopted to the last variant of texture development as a means for requiring minimal energy expenditure; this variant of deformation development explained details concerning strain hardening and resulting recrystallization in the considered samples.
The opinion is widespread in literature that by recrystallization of α-Zr, the main mode of texture changes consists of the rotation of the lattice by 30° about the basal axis [13, 15]. Some studies, however, do not support this opinion [16-18]. In the current chapter, an attempt was made to explain the discrepancies in experimental data concerning this question. In Fig. 13 PF{11
This rotation means that in the rolled matrix there are recrystallization nuclei situated at the boundary between two deformed grains with close orientations, have a middle orientation and swallow grains of the matrix due to growing into these neighboring grains at equal rates. This mechanism can work only when the rolled matrix has developed a deformation texture, so that the rotation of prismatic axes by 30o is not the only possible situation in some fragments of the sample; instead, a more complex process takes place in a coordinated manner within the entire matrix. The misorientation of some newly originated nuclei by 30o with the deformed matrix can promote its growth, but it does not create the recrystallization texture in the absence of the developed rolling texture.
As for the rolled single crystal N 3 (Fig. 13e, f), this case deserves separate consideration. Due to the position of its basal axis near TD (Fig. 5j), the largest part of this single crystal deforms at first by means of a prismatic slip and only by ε ≈ 7% the twinning is noticeable, overthrowing a portion of basal axes at the central segment of diameter ND-RD (Fig. 5m). PF{11
Recrystallization in cold rolled α–Zr single crystals N 2 (ε~30%):
Recrystallization in cold rolled α−Zr single crystals N 2 (ε~70%) :
Recrystallization in cold rolled α−Zr single crystals N 3 (ε~30% (
Recrystallization in cold rolled α−Zr single crystals N 3 (ε~30% (
Texture changes through the recrystallization of rolled α-Zr single crystals N 1, ε=80% (
The consideration of data regarding the recrystallization of a rolled α-Zr single crystal can be used for explaining the features of polycrystalline α-Zr sheets observed during their recrystallization. In Figs. 14 and 15, the results of X-ray diffractometric studies of iodide coarse-grained α-Zr rolled, respectively, by 50% and 70% and then recrystallized, are presented. In Figs. 16 and 17, analogous data are presented for α-Zr in a polycrystalline Zr-1%Nb alloy, rolled by 70% along and across the initial RD, respectively, and finally, recrystallized. In all Figures PF(0001), rolled samples are shown as (a), PF(0001) for recrystallized samples – as (b); GPF0002– as (c); subtraction diagrams S(ψ, φ) - as (d) and diagrams of correlation PF(0001) - GPF β0002- as (e). In Fig. 18, polycrystalline materials PF{11
According to models of texture formation in α-Zr [13, 19], the pole density in the central region of PF(0001) is conditioned by the input of grains experiencing twinning by deformation. Subtraction diagrams show that under recrystallization in the corresponding regions of PF for the deformed samples, an increase of pole density was absent. The reasons for this effect in polycrystalline α-Zr sheets are the same as for rolled single crystals. The absence of growth of α-Zr grains with orientations corresponding to the central region of PF(0001) has been noted in previous studies dealing with the recrystallization of products from Zr-based alloys [8, 20]. However, only experiments with the rolling and recrystallization of α-Zr single crystals have yielded a direct demonstration of the effect as influenced by twinning on the strain hardening and the subsequent recrystallization of α-Zr crystallites.
Recrystallization of polycrystalline rolled iodide α-Zr (ε~50%):
Recrystallization of polycrystalline rolled iodide α-Zr (ε~80%):
Recrystallization of Zr-1%Nb alloy rolled along initial RD (ε~50%):
Recrystallization of Zr-1%Nb alloy rolled across initial RD (ε~50%):
Texture changes through recrystallization of rolled Zr single crystals N 1 (a, b), N 2 (
PF (0001) (a) and PF{11
PFs (0001) (a) and {11
Shifts in the maxima rolling textures of polycrystalline Zr sheets by recrystallization are explained by the preferential growth of grains with basal axes orientations at the outward edges of these maxima. Here, there are more recrystallization nuclei and new grains that arise here absorb deformed grains localized within previous maxima textures (Fig. 14-17d). This shift occurs due to the inhomogeneity of strain hardening in the rolled samples, consisting first of all in its strengthening with increase of an angular distance of crystallite from texture maximum. This effect is clearly seen in Figs. 14-17 and has the same nature as bordering maxima textures with zones of increased strain hardening and intensive growth of recrystallized grains in rolled single crystals. The outward edge of different stable texture maximum slip systems, i.e., basal and prismatic systems, act together and since the volume of material to be deformed here is relatively small, simultaneous operation of both systems is possible. However, when moving inside the stable texture maximum, these slip systems begin to operate alternately in each point of the sample. As a result of recrystallization, new maxima grows at “slopes” of initial maxima in the rolling texture. Here, the strain hardening is sufficiently high for providing the preferential growth of arising nuclei, while at the same time the number of grains located at this segment of “slope” is sufficiently large for absorbing the entire deformed matrix. Obviously, this feature is typical for all metal material at the deformation stage and corresponds to the maxima arising in their deformation textures. Features of their recrystallization at the first approximation reflect the characteristics of strain hardening distribution, formed during the course of their rolling.
In terms of data related to the sample of Zr-1%Nb alloy rolled across initial RD, we observed that this sheet was characterized by the same features of strain hardening distribution and recrystallization development, the only difference being the space isotropy of constructed diagrams. As a result of symmetry, the obtained data were confirmed to be statistically safer.
In rolled Zr single crystal, deformed up to 80% by means of twinning and predominantly consisting of basal and pyramidal slips (Fig. 5b), after annealing at 580oC, a noticeable rotation of prismatic axes was not observed (Fig. 13). At the same time, in polycrystalline Zr rolled up to 80% (Fig. 15), after recrystallization, the sharp reorientation of prismatic axes became apparent (Fig. 18) by means of their rotation at 30o about the basal axis, coinciding with the correct texture maximum (Fig. 15b). It was evident that the rotation of prismatic axes about basal axes took place only in cases when recrystallized areas of α-Zr were deformed by rolling and with the inclusion of a prismatic slip, which according to Hobson’s scheme [2], develops within wide regions of stereographic projection adjacent to TD. The development of stable α-Zr rolling texture is always realized by active participation of a prismatic slip, since the existence of texture maxima at a horizontal diameter of PF(0001) is conditioned by the mutually balanced operation of a slip in the prismatic and basal planes [6]. Only in a case where texture maxima are distanced from ND by a comparatively small angular distance, as in PF(0001) of the single crystal rolled to ~50%, can one certify that a prismatic slip did not participate in the formation of texture.
Increasing the heating rate up to attainment of the recrystallization temperature (see Fig. 19c, d and f) or the presence of additional components in the texture of the rolled sheet, for example, T1+T2 (Fig. 20), prevents the rotation of prismatic axes.
Analysis of data regarding the rotation of prismatic axes under annealing of rolled α-Zr allows for ascertaining the mechanism of its recrystallization, established earlier by considering strain hardening distributions using GPF β0002. The presented data testify that in α-Zr by recrystallization, the mechanism is realized by combining two different aspects:
the rotation of basal axes in accordance with the strain hardening distribution, revealed with the help of GPF β0002 and predetermining the predominant growth of regions responding by their orientation to “slopes” of maxima in the rolling texture;
the nuclei of recrystallization formed in these regions during the course of their growth can cause (or not) 30о rotation of the matrix, depending on the participation of the prismatic slip in their deformation.
From nuclei arising in zones of increased strain hardening and having all possible orientations, the matrix chooses nuclei characterized by a 30°angle of misorientation at the boundaries as the best prospects for future growth. The increased heating rate by annealing of material and the presence in the deformed matrix of layers of grains with other orientations [8, 9] prevent such selection (Figs. 13, 20).
Dependence of the observed reorientation of prismatic axes on the heating rate of the sample can be explained in the following manner: the growth of recrystallization nuclei as part of the diffusion process requires a definite time period for the revealing of these nuclei, which are capable of swallowing the deformed matrix and to impose its own orientation as a result of the optimal angle of misorientation at grain boundaries.The presence in the deformed matrix of layers of grains with orientations of RD different from {10
Experiments with rolling and annealing of α-Zr single crystals helped to understand the mechanisms of α-Zr recrystallization in polycrystalline products, the absence of grain growth in regions of predominant twinning, rotation of the deformed α-Zr matrix about the basal axis and the conditions for the preferential growth of new grains.
Regions of the α-Zr matrix deformed by predominant participation of twinning were characterized by minimal strain hardening, by an insignificant number of recrystallization nuclei and a decreased trend towards grain growth, so that under recrystallization, these regions were swallowed by those that had higher strain hardening.
The necessary condition for 30o rotation of the deformed matrix about the basal axis through recrystallization of rolled α-Zr consisted predominantly of the participation of a prismatic slip in its plastic deformation.
In the course of α-Zr recrystallization the following mechanism for the preferential growth of definite regions was realized: these regions were characterized by relatively high strain hardening and orientations “slopes” of maxima in the rolling texture, where number of grains was sufficiently great in order to swallow the whole α-Zr matrix. Deformation of the α-Zr matrix at these “slopes” was realized by the simultaneous joint operation of basal, prismatic and pyramidal slip systems, whereas in central regions of maxima these systems operated alternately and provided minimal strain hardening.
By the quick heating of rolled samples with the purpose of their recrystallization, rotation of the deformed matrix by 30о about its basal axis did not take place, as in the case, when additional components were present in the rolling texture.
Surgical treatments of bone injuries patients in emergency departments worldwide each year due to involvement in rigorous athletic activities, social instability, traffic accidents, and prolonged human lifespan [1].
Bone defects, mainly induced by traumatic avulsions, sequelae of infection-induced bone sequestration, congenital malformations, or neoplastic resections, confront us with an extreme challenge for reconstructive surgery the need to induce bone regeneration to repair structural bone deficient [2] has inspired research on and development of a vast number of bone repair materials.
Diverse metallic materials are already established as biomaterials due to their high biocompatibility, low toxicity, and good strength–ductility relationship. Examples of these alloys are stainless steel (especially 316 L), cobalt and chromium (CoCr) alloys, and titanium (Ti) alloys [3, 4]. However, the low toxicity and mechanical properties of Ti alloys, specifically the elastic modulus, are more adequate for biomedical uses. From Ti alloys, the most common for dental and orthopedic applications are materials formed by Ti, aluminum (Al) and vanadium (V) like Ti-6Al-4 V and other β-phase-type alloys as the ones with high contents of β-stabilizers (V, Cr, molybdenum (Mo), Fe, niobium (Nb), and tantalum (Ta)) [5, 6, 7, 8, 9, 10, 11, 12]. However, several reports point to the V in the Ti-6Al-4 V as toxic [13, 14], being a motivation for exploring further V-free options. Moreover, the β-phase type Ti alloys have a good combination of mechanical properties and biocompatibility. Besides, the β-Ti alloys have a lower elastic modulus compared to other Ti alloys [15]. Considering that the elastic modulus is a key factor for the success or failure of the implant, this is a remarkable characteristic of these alloys [1]. However, the reported elastic modulus for β-Ti alloys ranges from 69 to 110 GPa [15, 16], being still far from that of human bone (lower than 30 GPa) [17].
To overcome this drawback, several Ti alloys are being developed and most of them are showing promising results in the matter of mechanical properties. A number of these metallic systems are being obtained through powder metallurgy methods to obtain functional porous structures. It has been widely reported that the porous surfaces assist on the fixations and ingrowth of organic tissue, improve the body fluid, reduce the mechanical mismatch due to lower elastic modulus values, and reduce the failure rate of implants [3]. Examples of the above are Ti and indium (In) as (Ti-In) [18], Ti-Mo [7, 8, 9], Ti, Nb and Tin (Sn) as Ti-Nb-Sn [10, 19], Ti and zirconium (Zr) as Ti-Zr [20], and Ti and silver (Ag) as Ti-Ag [21] alloys. However, some of the previous systems employ alloying elements that are still not widely studied, being a reason why several
Another route is the design and development of biomaterials based on widely explore elements as magnesium (Mg). This element has multiple advantages for biomaterials as non-toxicity, biocompatibility, biodegradability, increase strength of the bone, and has a low elastic modulus [3, 4, 22, 23, 24]. Low concentrations of Mg2+ play an important role in cells activity by stimulating the improvement of cell adhesion and extracellular matrix mineralization [25, 26]. Furthermore, Mg is the fourth most abundant element in the human body and is essential in digestion processes [22, 24]. The non-harmful degradation of an Mg, zinc (Zn) and manganese (Mn) as Mg2Zn0.2Mn alloy inside the human body has been demonstrated [23]. Based on the above, Mg is a feasible alloying element to boost the biocompatibility and possible control of biodegradability over the time of different biomaterials for medical purposes. The biodegradability of Mg can avoid the need for a second surgical process to remove the implant. The possibility to control such biodegradability is still under intense investigation [23, 27, 28]. Moreover, Mg is a potential alloying element to significantly reduce the elastic modulus. This could reduce the failure rate due to mechanical mismatch between the implant and the bone, and the occurrence of load shielding (absorption of mechanical stress by the implant) [3]. However, one of the main disadvantages of Mg as a biomaterial is that the degradation rate can be faster than the required to allow a complete regeneration of the organic tissue [29]. This is the motivation to explore the use of Mg as an alloying element instead of a matrix. Considering the already explained qualities of Ti biomaterials, it is a good candidate to join with the virtues of Mg.
Until now, few reports on Mg as an alloying element of Ti alloys have been reported [30, 31, 32, 33]. Deep research is still needed in the matter of optimizing Mg contents, processing parameters, and designing new systems that reduce the economic and health losses due to the failure of implants. The field of Ti-Mg alloys is emerging and is pointing as highly promising for biomedical purposes.
This book chapter contains the latest findings on the microstructural, mechanical, and biological properties of Ti alloys with Mg additions designed and obtained by the authors. Also, a description the most important techniques to obtain Ti-Mg alloys for biomedical application. An especial emphasis on the microstructure-properties relationship was made to assist on the guide for future efforts of the scientific community towards developing more efficient biomaterials.
As the largest dynamic biological tissue in the body, bones are composed of inorganic minerals as magnesium and metabolically active cells surrounded by a large volume of extra cellular matrix, and they form a rigid scaffold that has an irreplaceable role in maintaining life activities, including supporting the body and protecting visceral organs [1]. For bone repair, metallic materials are used to repair or replace the bone tissue damaged. The main materials used in orthopedics include stainless steel and Ti alloys because they are mechanically strong and resistant to fracture. However, there is a potential for the release of metallic ions and/or particles through corrosion and/or wear that trigger inflammatory responses that can reduce biocompatibility and lead to tissue loss [34].
Previously
Mg is the fourth most plentiful cation in the human body, and is an element essential in many metabolic processes, involved in the regulation of eukaryotic cell proliferation, structural functions are correlates with the enhancement of protein synthesis. Furthermore, is primarily stored in bone tissue, controlling growth of bone cells and accelerates the bone healing [36, 37], which has characteristics of bio-degradability, in the physiological environment can be eliminated and also the corrosion product of Ti-Mg alloys (Mg2+) does not cause unexpected complications because excessive Mg2+ are easily eliminated in the urine. Moreover, alloys fabricated using Mg elemental can present mechanical properties similar to those of bone, due to fabrication pores material, decreasing the elastic modulus [38, 39]. Once the bone resorption around stress-shielded are in bone fixation treatments is an important consideration within clinical sectors. Because of its versatility, metallic biomaterials based on Mg can contribute to biological properties and improve the osseointegration process [40].
Mg2+ is distributed in three major compartments of the body: ~65% in the mineral phase of bone, 34% in muscle, ~1% in plasma and interstitial fluid [41] and it has a radius of 0.65
Its charge density is approximately (.99/.65)3 3x more than that of ion calcium (Ca2+) and its affinity for electronegative ligands, almost always oxygen in biological systems, is much greater. Further, the Mg-oxygen bond length is approximately 2.05
Ti alloys can be fabricated using as alloying element Mg to create porosity, leading to formation of biocompatible scaffolds with lower elastic modulus by metallurgical and additive manufacturing process. Thus, the stiffness of Mg-based implants can be more easily tailored to match that of bone, which reduces the risk of stress shielding, a phenomenon that will be discussed in the next section.
One of the most conventional techniques to create porous metallic materials is by the powder metallurgy. Conventional manufacturing of Ti alloys via powder metallurgy involves: (1) blending of the powder to achieve a uniform particle size distribution and, if needed, mixing of the Ti powder with the required alloying elements; (2) shaping of the powder blend (this can be achieved via different manufacturing methods where the simplest and cheapest is cold uniaxial pressing); and (3) solid-state sintering (i.e. heat treatment at high temperature, below the melting point of Ti) generally performed under vacuum [46]. This technique is cost effective, allows for better control of powder size and introduction of desirable pores. When it comes to pore size and shape, these are related to the size of the starting powder, its shape, its size and the shape of the spacer used to promote porosity. Among all of them, Mg is a good candidate in the manufacture of Ti scaffolds, as its solubility in Ti is low. Furthermore, Mg2+ increases osteoconductivity [47] and does not present any type of biomedical inconvenience, such as toxicity.
The main works carried out using Mg are for Ti-Al-V alloys [48]. Wen et al., (2001) reported that Mg foam prepared with a porosity of 50% showed a compressive strength of 2.33 MPa, with Young’s modulus of 0.35 MPa, respectively [49]. Zhuang et al. [50] also evaluated the mechanical properties of porous Mg manufactured by powder metallurgy and the scaffolds with porosity from 36 to 55% showed a Young’s modulus value in the range of 3.6 to 18.1 GPa, closer to that of natural bone [50]. They also investigated the effect of porosity on biodegradation. In their study, it was reported that materials with greater porosity degraded more quickly, due to greater interconnectivity and surface exposure, conditions that maximize chemical reactions. Although Mg materials have a low elastic modulus, mechanical resistance and corrosion are limiting factors for their use [51]. However, the use of Mg together with other elements to form porous alloys such as Ti can be an interesting alternative. For the production of porosity using this technique, it is necessary to control parameters such as temperature and times of the sintering steps, in addition to the particle size of the powders, due to its commitment to mechanical properties [52]. Such variables strongly influence the morphology of the pores, that is, it can provide the same amount of porosity, but with different shapes and sizes. The mechanical properties are also affected, mainly those related to the ductility and dynamic properties of the material, as they depend on the porosity characteristics [53]. The pores attenuate the applied force, do not distribute it over a larger area and cause local stress accumulation, so that they can even serve as sites for crack nucleation [54]. The effect of porosity on mechanical properties depends mainly on the following factors: volume fraction of the pores and their interconnection, size, morphology and distribution. The most important parameters are the total porosity, the shape of the pores and/or contacts during sintering [55].
Porosity has a noticeable and well-recognized effect on mechanical properties. Porosity can increase stress concentration and cause fractures. It was demonstrated by Danninger et al. that this parameter is directly related to the mechanical properties of an alloy [56]. These factors can be controlled by adjusting the sintering parameters, compaction pressure and particle size [57]. Optionally, functional porosity can be introduced by adapting the particle size of the starting powders and the sintering conditions. In addition, powder metallurgy allows flexibility in alloy design, mixing pure Mg powders with different elemental or alloy powders. Due to the high affinity of Mg for oxygen, all handling of powders and samples, as well as subsequent sintering, must be carried out under a protective atmosphere of argon or under vacuum [58] conditions, residual pores can vary between 2% and 45%. When approaching porosities close to 45%, interconnection (percolation) of the pores appears.
In recent years, interest has increased in the application of additive manufacturing of Mg alloys for its biomedical application. This can allow the obtaining of complex shapes adapted to the patient, since it would be a personalized manufacture. However, and due to the physical properties of Mg and its alloys, the application of additive manufacturing by melting the alloy has not been easy, since the boiling temperature of magnesium is very low (~ 1091°C). Despite this, and due to the manifest interest in the biodegradation properties of Mg motivate its application as biodegradable implants [59], seeking the best combination between resistance to corrosion, wear, mechanical properties and biocompatibility [60].
Another factor influencing the scarce development of additive techniques in Mg alloys is the ease of obtaining coupling parts by injection casting processes or the extrusion capacity of these alloys despite the difficulty imposed by their hcp crystalline structure. In addition, the great reactivity with oxygen that Mg presents must be considered and also limits the application of rapid heating techniques that could cause the combustion of the metal. However, some of the technologies applied to other materials have been applied, by means of specialized teams that require work in protective atmospheres, which ensure the possibility of handling these alloys [61].
The potentially most interesting techniques for the manufacture of magnesium alloys are powder bed fusion (PBF), especially selective laser melting (SLM), widely used in the development of different Mg alloys [62, 63]. Powder bed fusion (PBF) is an AM process in which thermal energy is used to selectively fuse regions of a powder bed [64]. The powder bed contains metal, polymer, or ceramic powder as feedstock. An energy source directed towards the powder bed selectively scans and melts the top layer of the powder bed. The powder bed then lowers and a fresh layer of powder is spread over the melted layer. This process continues until the entire structure has been formed by stacking melted layers of powder.
Another way to use additive manufacturing related to magnesium alloys is to have a porous structure obtained by additive manufacturing as in the case of Perets et al. [65] who obtain the TI-6Al-4 V mesh by SLM and then infiltrate the Mg elemental into the holes. In this way they can obtain structures that do not collapse, although their resistance is not increased. The magnesium will present an accelerated corrosion by galvanic effect and depending on the size of the pores, an osseointegration will be available as it is a very biocompatible set.
Degradation capability of Mg gives a feature of bioactivity in bone formation that leading Balog et al. develop a bioactive metal system compound by structural material for dental implants, via extrusion from a powder mixture of Ti and Mg (4 and 12%) in low temperature [66]. Adding Mg, is possible to obtain a bioactive system and a decreasing of elastic modulus, further promote a good osteointegration due to the Mg resorption and the presence of pores where the bone ingrowth can be formed.
Mechanical properties in Ti parts that receive infiltration of Mg depends on the amount of Mg and matrix used. Studies published by Jiang et al. about infiltration of Mg in a scaffold of Ti-Mg (99,9%) was possible to control density from compaction pressure with volume fractions up to 60% Ti, which confers stiffness similar to those of cortical bone [67]. The use of non-degradable Ti matrices, as described in previous sections, is necessary as a non-degradable support due to its excellent biocompatibility, high resistance to corrosion and excellent mechanical properties. Similarly, efforts have been made to obtain porous Ti by means of spacer techniques [68] or by additive manufacturing processes such as selective laser melting (SLM) [69, 70]. Biodegradable Mg-based alloys are advantageous as fillers for bioactive implants because the release of Mg ions during corrosion
The Ti-alloys are already recognized as the most promising materials for dental and orthopedic implants. This is due to their excellent biocompatibility, good mechanical properties, superior corrosion resistance, and no allergic issues [72]. Adding Mg provides multiple biological advantages to Ti biomaterials. Some of these advantages are an elastic modulus closer to that of human bone, high biocompatibility, and low toxicity. The mismatch between elastic modules of human bone and dense metallic biomaterials is considered the main cause of implant failure. Typically, the elastic modulus of human bone has a maximum value of 30 GPa, while that for metallic biomaterials ranges from 100 GPa for chemically pure Ti to 230 GPa for Co-Cr alloy [15]. From the metallic biomaterials, the near β-Ti alloys have reported some of the lowest elastic modulus values ranging from 80 to 110 GPa [14]. Additionally, porous β-Ti alloys have obtained elastic modulus ranging from 43 to 75 GPa [4, 37, 73, 74], and it is expected that the addition of Mg will decrease even more such values [75]. An elastic modulus near to that of human bone improves the performance of multiple biomaterials used for dental and orthopedic implants. Considering that the microstructural effects of Mg additions are important to define the mechanical behavior of biomaterials, it is important to study the microstructure.
Due to the above, the effect of low (3 mass%) Mg additions into a Ti-34Nb-6Sn alloy was studied [28]. The selection of Nb and Sn was done for its β-stabilizing effect on Ti alloys [9, 76]. The β-phase has shown better mechanical compatibility with the human bone in comparison to the α-phase. This is due to lower elastic modulus values and a high strength-to-weight ratio [13, 14]. An elastic modulus of the implant that is close to that of the human bone decreases the mismatch of mechanical stress through the interface and avoids the damage of the organic tissue cells. Based on the above, the closer the elastic modulus between both, bone and implant material, the lower the probability of crack nucleation and failure of the implant [72]. Furthermore, a high strength-to-weight ratio allows reducing the thickness of biomedical metallic implants. Besides, the selected contents of Nb and Sn also contribute to obtaining elastic modulus of the Ti alloy of ~60 GPa [9, 74, 77]. The addition of 6 wt.% of Sn into the Ti-Nb alloy showed a good combination of corrosion resistance, strength, hardness, and lower elastic modulus [4]. On the other hand, considering the low elastic modulus of Mg (from 39 to 46 GPa) [78], it was added to reduce the elastic modulus of the Ti alloy. Besides, Mg is a natural component of human bones and is a required element for the metabolism process, i.e., Mg exhibits great biocompatibility, non-toxicity, and can stimulate hard tissue recovery [2]. The reported biodegradability of Mg2+ is one more of its valuable advantages [79]. The possibilities of its use in dental and orthopedic implants can be highly beneficial from controlling its degradation rate and ensuring its mechanical integrity during desire clinical periods. Moreover, the corrosion products of Mg, trigger the osteoconductivity of the bone [75]. For the last, the β-phase Ti alloys have shown a superior electrochemical performance that provides better resistance in corrosive environments as oral or body fluids [13, 32]. This is due to the surface TiO2 passive film that inherently protects these alloys [10, 32].
From the above-mentioned selection of components, four Ti alloys were prepared by powder metallurgy method. A typical four-stage route of milling – mixing - compaction – sintering was used. For this, measured amounts of titanium hydride (TiH), niobium hydride (NbH), and atomized Sn were used to obtain a mass ratio of Ti, Nb, and Sn of 60:34:6. The powders were mixed in a planetary ball mill and grounded at 200 rpm for 40 min. Posteriorly, the mixed powders were dried under a vacuum. Details of processing parameters can be found in previous work [28]. Half of the mixed powder was saved with the abovementioned chemical composition, while the other half was mixed with Mg powder in a 3 mass%. All the dried powders, whether with or without Mg, were compacted at 200 MPa for 15 s. To evaluate the microstructural and mechanical effect of Mg addition, two different sintering temperatures were used, 900°C and 1100°C. The sintering was carried out in a high vacuum resistive furnace for 2 h. A scheme of the elaboration process is shown in Figure 1.
Representation of the methodology to elaborate the Ti-34Nb-6Sn and Ti-34Nb-6Sn/Mg alloys [
Finally, two Ti-34Nb-6Sn (TNS) and two Ti-34Nb-6Sn/Mg (TNS/M) alloys were produced. The identification of the obtained alloys indicates the sintering temperature as postfix: TNS900, TNS1100, TNS/M900, and TNS/M1100.
For microstructural analysis, the samples were subjected to conventional metallographic preparation until a mirror-like surface. Final polishing with oxide polishing suspension (OPS) solution and hydrogen peroxide (10:2) was applied. Rietveld refinement of X-ray diffraction measurements was carried out for quantifying the present phases and estimating the lattice parameters. A Bruker/D2Phaser with Cu-Kα radiation was used at 30 kV and 10 mA. The measured 2θ range was 20 and 90° with a step size of 0.02° every 10 s. The Rietveld refinement was carried out by MAUD software (version 2.94) [80]. The morphology and chemical distribution of phases were studied by field emission scanning electron microscopy (FESEM) (ZEISS-ULTRA 55) and an energy-dispersive X-ray spectroscopy detector (EDS) (Oxford Instruments Ltda.).
Considering that the mechanical properties play a determining role in the performance of biomaterials, the elastic modulus was estimated by impulse excitation technique (ATCP, Sonelastic®). Hardness measurements were obtained using a load of 147 N by the Rockwell method (BECLA), using a spherical steel indenter with a diameter of 0.16 cm. More details of the whole methodology can be found in previous works [28, 29, 30, 31].
As result, the four alloys resulted in tri-phasic microstructures of α-Ti (under hexagonal compact (hcp), structure), β-Ti (under body centered cubic (bcc) structure), and segregation of Nb. The microstructures can be observed in Figure 2, where the light gray color corresponds to the matrix of β-phase, the dark gray to the α-phase, and the bright particles to the Nb segregation. As it was expected from the increment of temperature, bigger grain sizes can be observed in the samples sintered at 1100°C (Figure 2c and d) compared to those sintered at 900°C. Both, α and β-phases, are randomly distributed in the microstructure, however, the linear chemical composition through the microstructure is not homogeneous. This is due to lower contents of Nb and Sn, especially Nb, in the α-phase (Figure 2e). Due to the well-known β-stabilizer nature of both alloying elements [1, 3, 4, 5, 6, 7, 8, 9, 10], those chemical gradients were expected. On the other hand, the Nb segregation occurrence was reduced with the sintering temperature (Figure 2c and d), which indicates a better Nb diffusion in the matrix when temperature increases. However, the Nb particles are continuously observed at both sintered temperatures. From the Ti-Nb diagram phase, Nb has low solubility in Ti, so the continuous presence of Nb segregates was expected [3].
Microstructure of a) TNS900, b) TNS/M900, c) TNS1100, and d) TNS/M1100, as well as e) linear chemical gradients through α and β phases representative of the four studied samples. Adapted from [
The phases percentages estimated by Rietveld refinement from XRD measurements and the total porosity obtained by Archimedes method for the four alloys are presented in Table 1. The TNS900, TNS1100, and TNS/M1100 samples showed similar phases percentages. However, the TNS/M900 showed a reduced β to α transformation during sintering. Additionally, both TNS/M samples obtained higher porosity percentages in comparison with the Mg-free alloys. This was a clear suggestion about the reduction of diffusional processes when Mg is added into the Ti-Nb-Sn system. Thus, Mg addition has an apparent α-phase stabilization effect.
Sample | α (mass%) | β (mass%) | Total porosity (%) |
---|---|---|---|
TNS900 | 20.1 ± 0.3 | 79.9 ± 0.0 | 22 ± 1 |
TNS1100 | 23.8 ± 0.5 | 76.1 ± 0.0 | 11 ± 1 |
TNS/M900 | 35.3 ± 0.0 | 64.7 ± 0.0 | 29 ± 1 |
TNS/M1100 | 22.4 ± 0.0 | 77.6 ± 1.4 | 20 ± 0.5 |
Phases percentages and total porosity of TNS and TNS/M samples sintered at 900°C and 1100°C.
Furthermore, the increment of porosity with the Mg content could be related to the highest content of oxygen from the intrinsic passivation layer of Mg2+. When temperatures increase, the release of gas also increases, generating pores at the microstructure [81]. Additionally, the Mg powders acted as a spacer in the TNS/M samples. Comparing the low melting point of Mg (~650°C) with that of Ti (~1668°C), it is evident that a fraction of Mg is evaporated during sintering, while the Ti content remains constant. The partial evaporation of Mg assisted in the formation of the pores during sintering. It is well known that the porosity tends to decrease at higher sintering temperatures during powder metallurgy methods [82]. Thus, the reduced porosity for the samples sintered at 1100°C compared to the samples sintered at 900°C was an expected result. The porosity could influence the mechanical properties of the alloy, especially in the strength and elastic modulus. The mechanical properties will be discussed in the next Section 3.
Posteriorly, the previous Ti-34Nb-6Sn and Ti-34Nb-6Sn/Mg alloys were also reported through the same powder metallurgy methodology, except for sintering temperatures of 700°C and 800°C [29, 30, 31]. These samples will be identified as TNS700, TNS/M700, TNS800, and TNS/M800 for this book chapter. Compared to the previous TNS/M900, and TNS/M1100, the same three constitutive phases, α, β and Nb segregation, were observed at the TNS and TNS/M alloys sintered at 700°C and 800°C. For comparison purposes, representative EDS measurements of TNS/M800 and TNS/M900 are shown in Figure 3. Figure 3a is representative of the distribution of the elements in the samples sintered at 700°C and 800°C, while Figure 3b represents the distribution of the elements in samples sintered at 900°C and 1100°C. Similar distributions of the alloying elements were observed in both cases, except for the Nb segregates. Figure 3a shows a greater presence of Nb segregates in comparison with that representative of sintering above 900°C (Figure 3b). This can be explained by the lower solubility of Nb in Ti below 1100 K (~827°C) [3]. The solubility, together with the effect of temperature, could also be related to the smaller grain size in the samples sintered below 900°C (Figures 2 and 3). While the β-phase matrix of the TNS1100 has an average grain size of 15 μm, that of the TNS700 has an average of 6 μm. This could be due to the lower solubility of Nb at lower sintering temperatures generated more segregated particles through the microstructure. Those particles could act as a pin for grain growth. The pin-like behavior of Nb segregates in a Ti matrix has been reported before [16].
Comparison between alloying elements distribution of the Ti-34Nb-6Sn/Mg alloy representative of a) the sintered at 700°C and 800°C and b) the sintered at 900°C and 1100°C. adapted from [
From Table 2, the TNS700, TNS/M700, TNS800, and TNS/M800 samples presented higher porosity percentages in comparison with the samples sintered at 900°C and 1100°C (Table 1). Besides, the samples sintered at 800°C were more compacted than the samples sintered at 700°C. This was a confirmation about the higher the sintering temperature, the lower the porosity percentage. It is well-known that increasing the sintering temperature favors the diffusional processes and more compacted microstructures with higher relative densities, i.e., lower porosity percentages [82]. Besides, Table 1 and Table 2 showed an increment of porosity with the Mg additions for samples sintered at the same temperature. As it was explained before, the partial evaporation of Mg and its passivating oxide, contributed to the increment of porosity. Among the most studied spacers for powder metallurgy methods are carbamide, sodium chloride, ammonium hydrogen carbonate, and Mg [83]. From those, Mg has shown superior advantages over the organic spacers due to its good biocompatibility and good mechanical properties [79]. The increment of pore formation with the Mg additions could be beneficial for the implant by decreasing the elastic modulus. This topic will be covered in Section 3.
Sample | α (mass%) | β (mass%) | Total porosity (%) |
---|---|---|---|
TNS700 | 45 ± 0.3 | 55 ± 0.3 | 23 ± 0.5 |
TNS800 | 31 ± 0.2 | 69 ± 0.2 | 21 ± 1.4 |
TNS/M700 | 41 ± 0.1 | 59 ± 0.1 | 38 ± 0.6 |
TNS/M800 | 33 ± 0.3 | 67 ± 0.3 | 28 ± 0.3 |
Phases percentages and total porosity of TNS and TNS/M samples sintered at 700 °C and 800 °C.
Besides, from Figure 3, a larger pore size for the samples sintered below 800°C is notable in comparison with that for the samples sintered above 900°C. Macropores of ~100 μm in average were acquired (Figure 3a) for the samples sintered below 800°C. This was contrasting with the pores in the range from 5 to 35 μm obtained in the samples sintered above 900°C. This could be related to lower atomic diffusion resulting from the lower sintering temperature that reduced the bonding ratio in the samples. Porosity improves the bonding between bone and implant material, encouraging the anchorage and growth of the organic tissue [15, 75]. Thus, porous materials facilitate tissue generation enabling body fluid transmission [84]. Considering that allowing successful osseointegration is one of the main requirements of dental and orthopedic implant materials, the porous structures are highly promising for those applications. It has also been reported that macro-pores are beneficial for multiple biological processes as cell attachment, ingrowth of osteoblasts, vascularization, and osteoconductivity [85]. However, an adequate vascularization requires pores with diameters larger than 100 μm, specifically in the range from 100 to 500 μm [15, 86]. From the above and the fact that increments of porosity percentage used to assist on the decrement of elastic modulus [75], the alloys sintered at temperatures below 800°C could be more appropriate for dental or orthopedic implant applications. However, other mechanical properties as strength and hardness are also crucial for the performance of metallic implants. The mechanical behavior will be described in the following Section 5. This will contribute to clarifying the current concerns about the role of porosity
It is well known that the microstructural features have a strong effect on the mechanical behavior of metallic components. In this section, the mechanical performance of the Ti-34Nb-6Sn and Ti-34Nb-6Sn/Mg alloys will be explained based on their microstructural features described in Section 4.
As it was explained before, the Mg additions into dental or orthopedic biomaterials should assist the adequate biocompatibility and a good mechanical strength-elastic modulus relationship. Lower elastic modulus is expected with Mg additions. Moreover, lower elastic moduli are expected from higher porosity percentages, this is, for lower sintering temperatures. Simultaneously, higher hardness values can be expected from the more compacted samples, i.e., the less porous. To evaluate these hypotheses, hardness and elastic modulus measurements were carried out.
From previous works [28, 29, 30], Table 3 presents the Vickers hardness as a function of sintering temperature for the TNS and TNS/M samples. A general tendency to increase hardness as a function of sintering temperature was observed. This is congruent with the lower total porosity percentage in the samples sintered at higher temperatures. As it was explained in Section 2, the temperature increment fabrication atomic diffusional processes that result in bonding improvement and density increment. To observe the effect of temperature on the total porosity and hardness of the sintered alloys, Figure 4 compares those three values. A general decrement of porosity with the increment of temperature can be observed in the TNS (Figure 4a) and TNS/M (Figure 4b) samples. Simultaneously, the higher density of the samples resulted in higher hardness values for all the samples.
Sample | Hardness (HV) |
---|---|
TNS700 | 146 ± 16 |
TNS800 | 153 ± 12 |
TNS900 | 309 ± 44 |
TNS1100 | 411 ± 25 |
TNS/M700 | 92 ± 10 |
TNS/M800 | 120 ± 20 |
TNS/M900 | 226 ± 85 |
TNS/M1100 | 344 ± 24 |
Comparison of Vickers hardness (HV) between the TNS and TNS/M samples sintered at 700 °C, 800 °C, 900 °C, and 1100 °C.
Hardness and total porosity as a function of sintering temperature for the a) TNS and b) TNS/M samples.
For comparison purposes, the hardness of elemental Ti ranges from 1.3 to 2.0 GPa [16], while the hardness in the TNS and TNS/M samples ranges from 0.9 to 4.0 GPa. Besides, the minimum recommended hardness value for metallic biomaterials is 1.2 GPa for avoiding high wear damage susceptibility during chewing and daily oral processes [87]. However, the hardness of the natural human teeth ranges from 2.2 to 3.9 GPa [88]. Hardness values near to the ones of natural teeth could assist in decreasing wearing between teeth and implant. From the sintered TNS and TNS/M alloys, the TNS700, TNS800, TNS/M700, and TNS/M800 have hardness values below 1.5 GPa, this is, below the minimum acceptable for avoiding wear damage and being within the range of the hardness of natural teeth. As result, those four alloys cannot serve as a feasible biomaterial for dental applications. However, the hardness of the human bone ranges from ~0.3 to ~0.75 GPa [89, 90]. This means that all the sintered samples are within the acceptable hardness range for being applied as orthopedic implants.
From Figure 4b, it is also possible to observe a slower decrement rate of total porosity with the sintering temperature in comparison with that for the samples free of Mg additions (Figure 4a). This could be related to the abovementioned effect of Mg as a spacer. The partial evaporation of Mg created additional pores compared to the created in the Mg-free samples (TNS). As result, the hardness increment with sintering temperature in Mg-added (TNS/M) alloys also showed a slower rate of increment in comparison with the TNS samples.
As it was discussed in previous Section 4, the elastic modulus plays a key role in the success rate of dental and orthopedic implants. Considering that the elastic modulus is a measure of the stiffness of the material, it determines the resistance to deform a material in the elastic range. For increasing the feasibility of the implant, the constituent alloy should have an elastic modulus near to that of the human bone. The elastic modulus of human bone ranges from 5 to 30 GPa [15, 91].
For evaluating the elastic modulus of the sintered alloys, Figure 5 presents a comparison between the values obtained for the TNS and TNS/M samples sintered at different temperatures. Considering that the samples sintered at 700°C and 800°C were discarded as potential materials for dental implant applications, the samples TNS700, TNS800, TNS/M700, and TNS/M800 were not included in Figure 5. Lower elastic modulus can be observed in the samples with Mg addition compared to the TNS systems. This result was congruent with the lower hardness values of the TNS/M samples (Figure 4), which implies that these alloys are softer than the TNS for similar sintering conditions. As it was explained before, the lower hardness in the TNS/M samples resulted from the spacer-like behavior of the Mg powders. Besides, a tendency to increase the elastic modulus with the sintering temperature was observed. Being congruent with lower porosity percentages measured in the TNS1100 and TNS/M1100 samples.
Elastic modulus as function of temperature for samples sintered at 900°C and 1100°C with (TNSM) and without (TNS) Mg addition [
Comparing the obtained elastic moduli in the studied samples with the ones reported for human bone, the TNS/M900 sample can be the most adequate for its use in dental or orthopedic implants. Besides, the TNS/M900 alloy joins an acceptable hardness for biomedical implants and has adequate porosity features for triggering the anchorage between organic tissue and implant material. This is, the TNS/M900 sample combined the best microstructural and mechanical properties to be a potential biomaterial. However, the performance of these alloys under in vivo environments should be described to determine the feasibility of the studied alloys for biomedical purposes.
Development of biomaterials needs to focus on the biointerface construction to match the structure of the host tissue and to meet the mechanical requirements of specific tissue. In order to do that, metallurgic and additive manufacturing techniques present great potential in the development of Ti-Mg alloys, with complex shape formed by pores to be more biocompatible. It is critical to manipulate the surface by physical and chemical parameters to achieve the clinical purpose of the biomaterial, leading to a fast integration to the bone, due to the stimulating biological functions. In this review was showed that β -Ti-Nb-Sn alloy can be fabricated using Mg to create high content of porosity (20–38%), with an elastic modulus between 31 and 49 GPa close to bone tissue, and hardness close to commercial materials and higher than different parts of skeleton.
In this sense continuous studies and researches in this field is of great relevance for materials applied in life sciences.
This work was supported by the São Paulo State Research Support Foundation (FAPESP) [grants: 2017/13876-2; 2019/24237-6], by the Ministerio Español de Ciencia, Innovación y Universidades with Grant RTI2018-097810-B-I00 and the São Paulo State Institute for Technological Research, in the development of materials, whom the authors thank.
The authors declare no conflict of interest.
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