Introduction
The current trends in aerospace, automotive and bio-medical industries
involve the development of optimized light-weight structures providing
high strength as close as possible to that of conventional bulky
structures1. Also, foam based cellular materials have
been extensively used in energy absorption applications. However, one of
the limitations to use these structures is the poor availability of
control on the parameters such as pore size and wall thickness which
determine their properties. The development of lattice based cellular
materials with repetitive unit cells has been helping to tailor the
mechanical properties based on the loading requirements. Cellular
materials are termed as true cellular materials when their relative
density is less than 0.32.
Cellular materials become a rising star in the biomedical industry and
are a prospective replacement for fully solid implants. The use of solid
implants which have higher stiffness compared to the surrounding bone
results in stress shielding phenomenon, which might cause bone
resorption. This eventually results in implants loosening which requires
implant replacement3,4. The use of porous cellular
materials helps to adapt the implants stiffness to that of the
surrounding bone. Also, the presence of pores in the structure improve
bone regeneration which provides better fixation5,6.
Metal alloys being the first choice for implants, Titanium alloys such
as Ti6Al4V have been widely used in implants manufacturing due to their
high strength, corrosion resistance and
bio-compatibility7–9.
Additive manufacturing (AM) technologies such as Laser powder bed fusion
(LPBF) facilitates the production of intricate complex structures in a
short period of time. The same technique is employed in the current
study. LPBF is widely used in the production of cellular materials to
obtain higher precision compared to other AM
techniques5,10. The LPBF fabricated parts have certain
issues such as quite low surface finish, internal porosity, geometrical
deviation, residual stress and brittle material phases, which negatively
affectthe mechanical properties of the structures. Various studies have
concluded that the above-mentioned issues are highly influenced by
process parameters such as laser power, hatching distance, scanning
speeds and building direction11–16. Furthermore, heat
treatment processes such as stress relief and hot-isostatic pressing
significantly improves the mechanical performance by transforming the
microstructure to a more stable ductile α+β phase with the additional
advantage of eliminating residual stresses and internal
porosity17–21.
The mechanical properties of cellular materials are mainly characterized
by material type, cell topology, and relative density. Their
characterization is generally carried out through static and fatigue
compression tests. Cheng et al.22 compared compressive
properties of foam based and lattice based cellular structures with
different relative densities, indicating that the lattice based
structures had higher specific strength. Additionally, various studies
focusing on the influence of different types of unit cell and relative
density on the compressive behavior of cellular materials were carried
out23–30. It was reported that the strength and
stiffness of the structures enhances with increasing the relative
density in well agreement with Gibson-Ashby law2. This
increase depended on the morphology of the unit cell as well. Depending
on the local loading conditions in the struts, the unit cell topologies
can be grouped into two different categories of bending dominated and
stretching dominated structures. Stretching dominated structures are
characterized by high strength and stiffness since the struts are
subjected to axial loading conditions. On the other hand, bending
dominated structures do not possess any struts along the loading
direction and hence fail due to bending loads in the struts. Bending
dominated structures are more compliant compared to stretching dominated
structures which failed mainly due to buckling. The compressive behavior
of bending dominated structures consist of three regions, elastic
region, flat plateau region where strain increases with constant stress
followed by densification. However, in stretching dominated structures,
the plateau region consists of oscillating stress followed complete
densification of the structure despite cell
morphology29–34.
Cellular materials used in biomedical applications as well as aerospace
applications are exposed to cyclic loads. Therefore, understanding the
fatigue behavior of these materials has gained strategic importance in
the recent years. Various studies have indicated that the
compression-compression fatigue properties are dependent on various
parameters such as cell topology, stress ratio (R-ratio), heat treatment
and the presence manufacturing defects from AM
process20,29,35–42. Zhao et al.36and Yavari et al.31 have investigated the effect of
cell topology and porosity on the compression-compression fatigue
behavior. The studies have clearly shown that if the deformation in the
structure is bending dominated, plastic strain is progressively
accumulated, leading to the final fatigue failure. While the fatigue
crack growth is decelerated in structures that fail due to buckling. The
S-N curves normalized with respect to the yield stress indicate that a
single power law is followed by a particular cell topology despite the
difference in porosity. The effect of fatigue the stress R-ratio has
been studied using diamond unit cells. Specifically, it was found that
the higher the stress R-ratio the lower the fatigue strength at a given
number of cycles to failure43. The effect of post
manufacturing thermal and chemical treatments on fatigue properties were
also studied. Yuan et al.44 used two heat treatment
temperatures (750ºC and 950ºC) and showed that samples treated at 950ºC
had a broader plateau under static loading indicating better
performances under plastic strains. The fatigue endurance ratio was
increased by 0.5 – 0.6 times with heat treatment at 950ºC.
Hot-Isostatic pressing (HIP) and Chemical Etching (CE) treatments have
improved the fatigue strength of cellular materials by increasing the
ductility, eliminating some internal defects and improving the surface
finish20. The fatigue failure of cellular materials
can be divided into strain accumulation, crack initiation and crack
propagation. The strain accumulation in the structures is mainly due to
cyclic ratcheting as indicated in various studies36.
The crack propagation takes place in two steps, a first propagation in
the struts followed by propagation through the unit
cells45. Applied stress level and surface defects such
as roughness, defects and waviness have a greater influence on the
fatigue properties compared to internal porosity in the crack initiation
stage. On the other hand, parameters such as material, microstructure
and internal porosity influence the crack propagation
phase38,46.
The porous material used for bone implants should have mechanical
properties in range of the human bone for better fixation. Therefore,
mechanical properties of the implants are of primary importance at the
initial stage of implant fixation; once the bone regeneration is
complete, higher fatigue life from the implant may not be
necessary35. Hence, it is necessary to study a variety
of structures with varying cell topologies, irregularities, and pore
shape to understand the material properties. Benedetti et
al.47 analyzed the compressive behavior of various
types of cellular materials in presence of different porosity levels.
That study has shown that the properties of the different analyzed
structures range between the two extremities of a cubic structure and a
cross shaped cellular structure. Also, despite the absence of vertical
struts, cross shaped samples had the highest strength for the given
stiffness. Therefore, the below mentioned seven different topologies are
considered in the present study to explore their suitability to be used
as an osteo-integrative coating for solid implants. Since this porous
coating is commonly placed in the contact region between bone and solid
implant, the resulting load is entirely compressive. Therefore, the
knowledge of monotonic properties and fatigue strength under compressive
loads is fundamental for their proper design. For this purpose, cellular
lattice specimens were manufactured via LPBF using the titanium alloy
Ti6Al4V. Their cell architecture includes three regular structures
((#1) Cubic, (#2) Star and (#3) Cross, three irregular structures
obtained by skewing the junction of regular structures ((#4) Cubic
irregular, (#5) Star irregular and (#6) cross irregular); a(#7)
trabecular consisting of random arrangement of struts mimicking
trabecular bone topology structure. The samples were analyzed for
assessing porosity and struts dimensions. Compression tests have been
carried out using monotonic and cyclic loading conditions to obtain the
strength and stiffness properties (under loading and unloading) of the
considered structures The specimens have been subjected to
compression-compression fatigue loading with an R-ratio of 0.1 and
loading between 0.1-0.8 yield load. One specimen from each topology has
been employed to visually detect the deformation pattern in the
different investigated structures. Fracture surface analysis has been
carried out to show the failure mechanisms under quasi-static loading
and cyclic loading. The effect of geometrical irregularity on the
fatigue performance has been explicitly shown in the normalized S-N
curves. The results of the regular and irregular topologies have been
carefully compared with the trabecular based topology for a better
understanding of their behavior with the future optic of possible
systematic employment in biomedical applications.