Titanium based bone implants production using laser powder bed fusion technology

manufacturing parameters and post-processing techniques affect the obtained microstructure leading to various mechanical, corrosion and biological behaviors of the manufactured titanium. All of the features are considered in the light of speciﬁcations and needs of bone implant applications. The most critical disadvantages of the L-PBF technology, such as residual stresses and leading deformations are introduced and the potential solutions are discussed. Moreover, the manufacturability of porous bone implants that causes beneﬁt and harm in L-PBF applications are assessed.


Introduction
Biomedical implants may replace damaged tissues, organs and joints [1].They require a homogeneous and stable microstructure to obtain integrity according to International Standards Organization (ISO) 20160 [2].In an ideal implant, proper functioning with no further need for revision surgeries should be possible as well [3].Most biomedical implants with an orthopedic purpose are used for maxillofacial, spinal, hip and knee replacements [4].The expected properties from the implants are close to the characteristics of the target bone in terms of adequate mechanical strength, modulus of elasticity and corrosion behavior as well as biocompatibility to avoid negative effects to humans [5].
Functional and biomechanical needs [6] determine design criteria and material choice for the bone implants.Various biomaterials are utilized including metals, ceramics, polymers, and composites to attain biocompatibility [7].Metallic implants, including stainless steel, cobalt-chromium and titanium alloys, are preferred for joint replacement and fracture fixation of bones when high mechanical reliability is required.Titanium is used due to its low density, excellent biocompatibility, good mechanical properties, high corrosion resistance, low allergic and chemical reaction rate in intra-body implants [8,9].The most researched and preferred titaniums are commercial pure titanium (cp-Ti) and titanium aluminum vanadium alloy (Ti6Al4V) [10].Ti6Al4V, which consists of 6% (wt.)Al and 4% (wt.)V, has a dual microstructure as aþb phases at room temperature while commercially-pure titanium (cp-Ti) has a lower mechanical performance than Ti6Al4V due to its single-phase (a) microstructure [11].On the other hand, cp-Ti is highly biocompatible as it does not contain Al and V which have toxic effects for the body [12].
An important disadvantage of titanium implants is the high elastic modulus leading to stress shielding effect causing bone resorption due to unconformity of elastic modulus between cortical bone (10e17 GPa) and titanium (105e114 GPa) [13,14].To address this major problem leading to revision surgeries and high discomfort for patients, similar elastic modulus to cortical bone can be attained with porous structured titanium implants [15,16].Moreover, osseointegration is enhanced by creating biomimetic pores with larger surface area.This approach provides a better antibacterial property and prevents infection [17].Implants with a complex geometry can be designed and produced with additive manufacturing (AM) due to its inherent nature of layered manufacturing [18,19].Advantages of the AM processes can pioneer optimum success in bone tissue engineering such as enabling new model approaches, capability of precisely reproducing patient-specific implants, in other words customisation, acceleration of healing and reducing implant rejection, adequate and customized mechanical properties [20,21].Among different AM technologies, laser powder bed fusion (L-PBF) is the most widely used process for producing titanium implants in need of enhanced mechanical properties and feature complexity [18].The production of porous titanium implants with L-PBF aims to decrease the elastic modulus while retaining biomechanical properties by integration of micro pores into biomedical implants.This review article focuses on L-PBF technology for manufacturing porous and solid biomedical implants from cp-Ti and Ti6Al4V.First of all, the effects of L-PBF process parameters and microstructure is reviewed on the mechanical, corrosion and biological properties of cp-Ti and Ti6Al4V biomedical implants.The disadvantages of L-PBF, such as surface defects and residual stresses, are covered among precuations to reduce those problems.Finally, the advantages obtained with hierarchical porosity as well as economic factors are reviewed in this study.

Additive manufacturing for biomedical implants
Additive manufacturing sequentially processes raw materials by depositing from an upper source or consolidating material in a basin or a bed.The technology is based on decomposing a part's 3D digital model into a set of flat geometries such as slices or layers, which are then built one on top of the other, adding the required material to form each specific layer.The recent developments in this group of processes have enabled biomimetic biomedical implants in the designed geometric shapes from preferred materials such as metal, ceramic, polymer and composites [22].
The AM process chain begins from a computer aided design (CAD) file until the final application in the targeted area as shown in Fig. 1.AM enables fabrication of replications of complex anatomical models for surgical simplifications as well as tailor-made implants with desired geometry.The patient 3D imaging data is taken with computer tomography (CT) scanning to catch accurate geometries of the matchless implants during the design process.Moreover, finite element analysis (FEA) on CAD design and mechanical experiments on as-built specimens can be preferred and compared to prove the mechanical strength of orthopedic implants [23e25].The CAD file is then converted to STL (Standard Triangulated Language) file format to generate triangular facets.Using special algorithms, it is sliced into two dimensional digital layers [26].The AM process continues with parameter set-up, manufacturing, cleaning and post-processing steps.
AM technologies are categorized under seven different groups based on the joining mechanisms according to ASTM 52900 as (a) material extrusion, (b) direct energy deposition (DED), (c) binder jetting, (d) sheet lamination, (e) vat photopolymerization, (f) material jetting and (g) powder bed fusion [27].In the PBF technologies, the powder on the platform is heated until the melting temperature by electron or laser beam energy source.After that, rapid solidification step starts and thin-grained microstructure forms.As this process is repeated layer by layer, the solid part exposes to the heating.The thermal cycle causes metastable microstructure and heat transfer to determine the melt pool expansion.Each powderbased AM technology shows different heat transfer capacity because of varied working principles [20,28].The most preferred heat sources are the laser and electron beams which are applied as L-PBF and E-PBF technologies for the biomedical implants [20].E-PBF technology provides an alternative method to L-PBF as more efficient production and affordable cost in vacuum instead of conventionally produced biocompatible metals which have hard workability and expensive cost [10,11].In the L-PBF process, the energy input from the laser affects the microstructure, the cooling rate and phase transformation differently for each shape of metal manufactured part, therefore solid and porous models are evaluated separately.Cooling rates and input energy levels show different reactions on the models.

Laser powder bed fusion (L-PBF)
In laser powder bed fusion which is also known as Selective Laser Melting (SLM), Direct Metal Laser Sintering (DMLS), Laser Cusing, Direct Metal Laser Melting, etc., a focused laser beam selectively scans a working plane on a powder bed, and then the powder fuses based on the CAD data of the determined slice [20].The L-PBF process and equipment schematic are shown in Fig. 2 [18,29e34].As depicted, the entire production is carried out in a build chamber loaded with a protective gas (nitrogen or argon) to prevent oxidation at high temperatures.
There are over hundreds of process parameters defined in the L-PBF process.These parameters need to be correctly set since they, as well as their interactions, have crucial influence on the part quality.The most influential and researched L-PBF parameters, namely laser power, scanning speed, layer thickness and hatch distance, have a prominent effect on the output characteristics [35].Laser power has more dominant effect on the process temperature changes compared with laser scanning speed [12].The combined effect of these parameters are given with linear (LED) or volumetric energy density (VED) equations as indicated as Eqs.( 1) and (2).At VED, generally expressed as J/mm 3 , equation, P symbolizes laser power, v describes the scan speed, t is the layer thickness and s is the hatch distance [36] while LED is described as J/mm, P/y [12].Although VED formula is the precursor path that helps to make surface defects and pores insignificant, the parameters of VED cause inconstancy between (real) efficient energy density and formulated VED [37e41].Complex melt pool physics is not explained because of ignoring interaction  among heat source and powder by VED formula, as a result pore variety at external window comes out instead of the presenting narrow range [38,40e42].Ferro et al. [38] stated that efficient new VED (VED Eff ) formula was generated to tackle this unreliability with the adding unaccounted parameters which are material property and absorbed energy factors.In Eq. ( 3), VED Eff (J/mm 3 ) gives that a is the material thermal diffusivity, b is absorption factor and F is the laser spot diameter (mm), the same reference article presents the narrow, low and more steady defect porosity range through VED Eff application than VED (Fig. 3).The sufficiency of VED is described that the maximum temperature in the melt pool touches to boiling temperature of the metal.Conduction or keyhole mode melting regime generates via input energy to the melt pool which determine the depth.If the max.Temperature via VED rises above the boiling point, a "keyhole mode" melting regime generates, which creates a fairly deep melt pool and remelts several layers and also causes melt pool vaporization and pore formation [42,43].If the max.Temperature is below the boiling point, the conduction melting regime creates a shallow melt pool.The reason for presented P=√V instead of P=V in Eq. ( 3) is that the effect of increasing laser power is more dominant than decreasing laser scanning speed on melting regimes [44].The results of Bertoli et al. [42] supported that the keyhole mode formation is observed between 300 and 500 W laser power range against 100 and 200 W. Another approach from King et al. [43] showed that 100 W and 100 mm/s reaches to 2306 C peak temperature while 7326 C is obtained with 500 W and 500 mm/s for stainless steel.

LED ¼
Laser Power ðPÞ Scan Speed ðVÞ (1) Various performance criteria are taken into account in optimizing L-PBF process parameters depending on density requirements and others such as material composition, residual stresses, productivity, surface quality, mechanical properties, and geometric accuracy.L-PBF process parameters are varied with raw material characteristics and minor differences in the machine components and control software.Raw material characteristics include powder particle size distribution, chemical composition, particle morphology, etc.While differences in machine components involve powder laying method or gas circulation dynamics [45e51].

3.
Biomedical implant applications with L-PBF manufactured Ti6Al4V and cp-Ti L-PBF manufactured titanium parts can be utilized to substitute bone structures in biomedical applications.The performance of those parts is therefore expected to be aligned with those yield by the cortical and cancellous bone tissues.The mechanical properties of fully dense titanium samples normally exceed those of the cortical and cancellous bone, which is an important point for consideration as shown in Table 1.The most prominent factor to determine the roles of bone implants in the body is whether the load-bearing or non-load bearing [52].While the superior strength value of Ti6Al4V is preferred in regions where titanium carries the load, such as hip, spinal or knee implants [53] (see Fig. 4), the lower strength of cp-Ti compared to Ti6Al4V has made it preferred as fixation mini plates and screws in non-load bearing cranio-and maxillofacial surgeries [54].In addition, cp-Ti fits better on the defective area without leaving any spaces due to its higher ductility against Ti6Al4V material structure [12].The elongation at break value which is 15e24% of cp-Ti in comparison to a value of 6e10 % for Ti6Al4V explains the ductility.The elastic modulus of cp-Ti (105 GPa) is slightly closer to the bone than Ti6Al4V (110e114 GPa), which may prevent the stress shielding effect, however still is so high.Nevertheless, this brings an interesting opportunity to design and configure the parts with a designed porosity or even as a functionally graded structure.

3.1.
Commercial pure titanium (cp-Ti) phase with body-centered cubic (bcc) forms [68].Three types of a are formed based on the cooling level and solute capacity during the cooling process after passing through the b phase.
The high cooling rate causes typically martensite phase alteration (a 0 ), medium cooling level causes the limited distance diffused alteration (massive conversion) and low cooling rate causes long distance diffused (a) phase transformation.Various morphological features are therefore seen according to the cooling rate.Lathlike martensite and massive martensite a 0 structures are attained when the cooling rate is over 500 C/s (see Fig. 5) as mentioned above.The microstructure of cp-Ti is predominantly plate-like a and acicular a in cp-Ti if the cooling rate is below 90 C/s, also, it is fine-wavy acicular and basket-weave a from 90 C/s to 500 C/s [12,69e72].
The iron content is also highly effective in cp-Ti phase transformation as well as the cooling rate.Iron contributes to massive alterations along with quick cooling.Moreover, adding iron delays the long-distance diffusional alteration at low cooling temperatures.The surface topography of a alloys forms dispersed b phase recrystallized a grains that presents each cp-Ti grade content has a certain amount of iron, which is a strong b stabilizer.Iron owns weaker dissolvability in a phase and therefore it does not allow the b phase improvement along with cooling or hardening.b phase stays steady in an ambient temperature.The amount of iron relies on the cp-Ti grade level and increases with raising grade count (grade 1, 2, 3, 4).Iron supports to manage the grain size along recrystallization.Minor grain dimensions can be improved to enhance yield strength if requested, as defined in the Hall-Petch relationship in Eq. ( 4).As the b phase becomes formable in cp-Ti, it can play a role as a crack prevention zone under most loading conditions [68].The yield strength is s, friction stress is s 0 , grain size is d, positive constant of yielding is k.

Material composition and microstructure
Ti6Al4V has a dual microstructure as aþb, a phase as hcp crystal form with b phase as a bcc form [11].  C/s and 410 C/s, refined a-lamellae is observed and causes mechanical properties improvement [75].The diffusionless phase transformation from b phase to a completely martensitic a 0 phase occurs instead of the aþb if the cooling rate above the 410 C/s [28,75].Each of mentioned phases, which are martensitic a 0 phases, the huge volume of Widmanst€ atten a-laths and slight b phases along a-laths, possess b christallographic grain boundaries [75,76].Martensitic a 0 phase is organized horizontally and regularly between b grain boundaries [77], however, Widmanst€ atten a-laths with b grains are similar to basket-weave microstructures.Refined and needle-shaped structures are seen for both martensitic a 0 and Widmanst€ atten a-laths.Although both of a structures show high strength properties, coarser Widmanst€ atten a-laths are thicker (<10 mm) that lead to defect formation when the specimens are tested under applied force as a result, the weaker strength is obtained [75,77].

The effect of microstructure on the mechanical properties of L-PBF manufactured titanium implants
The conventionally manufactured cp-Ti and Ti6Al4V alloys have lower mechanical performance than L-PBF fabricated cp-Ti and Ti6Al4V (Table 1) since grain refinement with the martensite a 0 phase of L-PBF support to provide superior hardness, compression and ultimate tensile strength (UTS) through higher cooling rate [71,78].In terms of stiffness and strength, this is preferable while for ductility, there is a loss.Balc et al. [79] stated that the microhardness value of L-PBF manufactured Ti6Al4V is found as 398 HV while the minimum medical desired value is 349 HV.The Vickers hardness values of L-PBF manufactured cp-Ti and Ti6Al4V are 261, 409 HV, respectively at other articles [15,80].Also, obtained higher mechanical properties via L-PBF cause to decrease the wear rate of cp-Ti (Fig. 6) [12,30,31,81].Gu et al. [82] reported that optimum L-PBF manufacturing parameters give a high hardness and reduced wear rate.
Al element is a powerful solid-solution developer in Ti6Al4V [83], thus less crack propagation occurs during tensile testing because the dislocation of Ti6Al4V can be hindered compared with L-PBF manufactured cp-Ti.Zhou et al. [78] stated that the tensile strength of cp-Ti and Ti6Al4V is 650e990 MPa and 1199e1334 MPa, respectively.Ti6Al4V shows therefore super high strength against the cp-Ti via L-PBF manufacturing [78,80].Nevertheless, mechanical advantages of L-PBF manufactured cp-Ti might help to destress of Ti6Al4V popularity using for biomedical implant because the mechanical strength of L-PBF fully-densed cp-Ti can higher than determined value range of cast Ti6Al4V (930 MPa) [12,30] in Table 1.
Investigating the optimum levels of main parameters (laser power and laser scan speed) for cp-Ti fabrication, Ataee et al. [12] stated that the laser power effect on the thermal gradient is higher than the scan speed.Thermal gradient plays an influential role in the cooling rate leading to significant microstructural changes.Thus, there is a significant correlation between microstructure and the selected laser power [84].On the other hand, increase in scan speed is directly proportional to the cooling rate and triggers therefore martensitic a 0 phase formation [71,82].The oxygen level is the another factor which affects the mechanical properties of the cp-Ti.Increased oxygen level of L-PBF building chamber during cp-Ti manufacturing might enhance mechanical strength and hardness due to martensitic a 0 phase [84] although at an expense of oxidation.Kwasniak et al. [85] stated that adding oxygen rises to strengthening of cp-Ti while retaining ductility.As oxygen contents of L-PBF up to 0.21% [12] and 0.4% [70] rise respectively, the mechanical properties improve as shown in Table 2.  plays a vital role in preventing toxic metal ions release to the body [78].

Effects of microstructure to corrosion features of L-PBF manufactured titanium implants
The corrosion behavior of L-PBF Ti6Al4V is weaker compared with L-PBF cp-Ti [3,80,86].Xiao et al. [87] assessed that L-PBF fabricated cp-Ti and Ti6Al4V corrosion resistance in Hank's solution for 72 h as a function of immersion test and then electrochemical parameters measurements which are corrosion potential (E corr ) and passivation current density (i corr ) according to Tafel curve.Cp-Ti current density is higher than Ti6Al4V i corr .Ti6Al4V pass from the passive layer to the active layer easily, while cp-Ti's passive layer is a wider passivation potential.Ti6Al4V is so vulnerable to Cl À ions.Thus, it is prone to the pitting appearances on the surface and presents weaker stability and low passivation to corrosion versus cp-Ti [87].Zhou et al. [78] stated that cp-Ti (3.4 Â 10 À7 A/cm 2 e 0.579 V) shows greater stability and performance of dense of passivation than Ti6Al4V (3.2 Â 10 À7 A/cm 2 e 0.639 V).Expected corrosion damage is thereby less, and a strong blocking effect of cp-Ti occurs against body fluid ions (see Fig. 7).Also, the reason why cp-Ti and Ti6Al4V show different corrosion resistance is that Al (a) and V (b) phase stabilizers.b phase is more constant according to a 0 or a phase while a 0 corrodes before the prior b phase.The b phase causes a denser passive film, but the oxide layer of cp-Ti as a 0 phase exhibits more stable corrosion resistance according to Ti6Al4V.The less b and more a 0 phases lead to instability on Ti6Al4V [87].Apart from the L-PBF Ti6Al4V and cp-Ti, also higher amount b phase resides in L-PBF Ti6Al4V according to conventional Ti6Al4V due to the rapid cooling rate of L-PBF, therefore conventional method can show less corrosive resistance in basic artificial saliva and deionized water mediums (pH > 6).On the other hand, acidic mediums (pH < 6) might lead to that L-PBF Ti6Al4V behaves more corrosive because of instability of passive layer which interacts with the acid, thus more a phase forms instead of b [88].Sharma et al. [89] stated that the various corrosive mediums (NaCl, NaOH, H 2 SO 4 ) and simulated body fluid (SBF) are used for the electrochemical corrosion experiments of cast and L-PBF Ti6Al4V.H 2 SO 4 acid behaved the most corrosive medium during the experiments, meanwhile cast Ti6Al4V exhibited more stable under each medium.

Effects of physicochemical surface characteristics on biological properties of titanium implants
The biocompatibility feature of Ti6Al4V implant manufacturing through L-PBF technology is crucial in terms of being preferred in the body instead of cast techniques.The cell viability of L-PBF manufactured Ti6Al4V shows high performance as much as commercial cp-Ti cast manufactured bone implant [90].The bone cell proliferation and bone marrow formation are great noticeably at L-PBF manufactured porous cp-Ti compared to Ti6Al4V in vivo experiments according to histologic outcomes [53] as shown in Fig. 8. Ti6Al4V is also found to be less biocompatible in comparison to cp-Ti because Al and V have cytotoxic effects on the body [54,91].Apart from the Ti material properties, also surface roughness is impactful on the cell activities [92].While L-PBF manufactured Ti6Al4V samples have 6e14 mm R a values [53], R a of L-PBF cp-Ti is between 9 and 10 mm [70] and 3.3e4.2mm [93], but the limit of R a is up to 10 mm according to ISO1442-88 for medical practice [53].Balc et al. [79] also indicated that the surface roughness of Ti6Al4V smaller than 4 mm with optimized L-PBF parameters.It is important to note that the surface roughness obtained in L-PBF is highly dependent on process parameters and vector types, such as upskin, downskin, contour, etc.
Interaction starts rapidly between the protein of the cells and the material at the upper layer after implantation.The proteins collect based on knowledge of the surface texture and chemical features about the material surface.The biological reaction is initiated by tissues and mainly by cells.The body response makes available cell attachment steadiness and stimulates cell growth and differentiation.Another crucial point for connection is the wettability which triggers the surface protein and cell adhesion [94].The wettability is recognized with hydrophobic and hydrophilic action of the surface.Cells might be damaged when they interact with both superior hydrophobic and hydrophilic surfaces.Thus, the balance is quite critical to achieve sufficient wettability on the surface [95].Also, the contact angle among material surface and cells, which changes the cell growth, is relevant to the wettability [94].A hydrophobic surface has 90 e150 of contact angle.The superhydrophobic layer is greater than 150 , the hydrophilic layer is between 90 and 10 , and the super hydrophilic surface is under the 10 (Fig. 9) [95].Menzies et al. [96] studied how the contact angle alters the biocompatibility of cp-Ti.The lower contact angle provides superior wettability and supports cell adhesion.

3.4.
Post-processing methods on L-PBF manufactured cp-Ti and Ti6Al4V alloys Biomedical implants with qualified surface quality can be obtained through post-processing applications [97] following L-PBF fabrication with optimized scan strategy [36].The postprocessing treatment methods aid additive manufactured metal bone implants under microscale to enhance their clinical application use with surface improvements that trigger higher cellular activities and achieve closer bone mechanical properties [98,99].The methods are hot isostatic pressing (HIP), sandblasting, chemical etching, vibratory finishing and polishing (PL).Hot isostatic pressing, which is the heat treatment under high pressure and temperature, reduces the defect porosity despite lower mechanical properties in comparison to as-fabricated AM metal parts.HIP promotes minimizing pores and improve fatigue resistance, at the same time, support cell activities improvement on the surface and transform microstructure.Furthermore, Liu and Shin [28] determined that HIP supports enhancing the ductility of Ti6Al4V L-PBF parts.Sandblasting and polishing are crucial applicable post-processing factors to partially clean un-melted particles from the surface and reduce surface roughness [28,100].Also, the sandblasting for cp-Ti provides superior cell adhesion and activity [101], however Yan et al. [102] stated that sandblasting leads to the crack formation on the Ti6Al4V surface.Chemical etching is another method generally applied with nitric acid (HNO 3 ) and hydrogen fluoride (HF) solutions.More homogeneous surface functionalization and higher penetration at internal surfaces can be obtained through this acidic content, even interconnected porous structures [100,103].
Jamshidi et al. [99] reported that the effects of HIP, HIP-PL, HIP-Sandblasting and HIP-Chemical etching applications of L-PBF Ti6Al4V parts are compared with each other.HIP process  caused to occur equilibrium aþb phase altering from martensitic a 0 phase.This phase transformation led to reducing mechanical strength while increasing ductility due to small b fractions among vast a-Ti phases.Although sandblasting and polishing applications with HIP process cleaned partially un-melted particles from the surface, the peaks and troughs still remained at the sub-layers and affect negatively tensile strength and fatigue resistance [102,104].The comparison of HIP-Chemical etching with the other methods for fatigue resistance under the cycle gave the best fatigue performance, which reduces surface roughness and automatically improves cellular activities [99].Pyka et al. [98] expressed that chemical etching and electrochemical polishing were applied together with HF acid to L-PBF Ti6Al4V to achieve homogeneous surface roughness.Partially un-melted particles on the surface were cleaned via the chemical etching method, meanwhile the surface roughness was reduced uniformly by electrochemical polishing.The penetrating of acidic solution to internal porous structure leads to diminish strut thickness.The mechanical properties, therefore, decrease after the process.It may be beneficial to consider the thinning of the part at the post-processing while designing the part.

4.
Challenges and opportunities of cp-Ti and Ti6Al4V biomedical implants production by L-PBF

4.1.
The drawbacks of L-PBF manufacturing process for titanium biomedical implants Extremely high-temperature gradients and rapid solidification lead to the uncontrollable development of off-balance phases and microstructures.The stability, configuration and thermal behavior of the melt pool along L-PBF can largely determine the occurrence of some specific problems, which are layer separation (delamination), distortions due to residual stresses and balling effect [110].Along with the stress accumulating in L-PBF-manufactured parts, a great tendency to shrink occurs during the liquid-solid conversion.The increasing (residue) stresses accumulated during cooling is one of the main reasons for the layer separation or distortions in the manufactured part [111].The unsteadiness of the melt pool formed by using the melting mechanism of the entire target area [15] is responsible for the lump in the liquid along with the molten cp-Ti.This situation is known as "balling effect," and this undesired effect can be eliminated with process parameter optimization (Fig. 10) [71,82].
The melt pool normally contains out of equilibrium phase transformations and numerous heat-mass transfers throughout the L-PBF process.L-PBF parameters (laser power, spot size, layer thickness, and scan speed) and the features of powder material (grain dimension and morphology) make a significant contribution in evaluating liquid surface tension and the status surface topography.Also the metallurgical structure of the melt pool and the mechanical features of the fabricated material are assessed with L-PBF parameters and powder features [30,82].
The reduction of voids and defects on the microstructure improves the fatigue life span [112,113].However, L-PBF sourced pores and fatigue cracks cause to weak fatigue resistance [114].The elimination of voids and defects via heat treatment improves the fatigue life [112].Apart from that, also ISO 20160 standard for implant material requirements indicates uniform and consistent surface topography guarantee to obtain durability and steadiness of biomaterials [115].High thermal stress, refined grains and strongly textured Tailor-made cp-Ti implant via L-PBF production.
[25] Finite Element Analysis (FEA) was utilized to determine position of screws for preventing implant loss in vivo and to provide stress analysis of mandibular screws for optimization.cp-Ti 210 30 1000 58 210 Mentioned that hydrogenerated-dehydrogenerated cp-Ti is more cost-effective than gas-atomized Titanium. [108] Obtained greater mechanical strength, lower corrosive behavior and higher biological properties of HDH Ti with L-PBF manufacturing.compared to gas atomized L-PBF Ti. cp-Ti (Grade 1) 80 50 571 28 140 Biocomposite coating was applied on surface of L-PBF manufactured cp-Ti and surface of sandblasted L-PBF cp-Ti parts. [101] The result was that sandblasting technique demonstrates superior cell adhesion.cp-Ti (Grade 1) 90 25 500 e 180 Acicular a ' martensite phase with SLM manufactured cp-Ti converted to equiaxed a phase after the heat treatment at 650 . [109] The equiaxed a phase is dominated by innate acicular a ' phase.
j o u r n a l o f m a t e r i a l s r e s e a r c h a n d t e c h n o l o g y 2 0 2 2 ; 1 7 : 1 4 0 8 e1 4 2 6 microstructure of L-PBF titanium samples are not convenient for biomedical implant material due to mentioned problems above (deformations, cracks, balling effect, surface defect porosity) and low ductility even manufactured according to standard [116].Although L-PBF manufactured titanium may have high UTS and low elongation break (ductility) due to fine acicular a ' martensitic microstructure, these mechanical values may not be the same at different directions leading to mechanical anisotropy as a result of L-PBF technology [109].
To improve elongation break and eliminate anisotropy of L-PBF manufactured titanium implants, the process needs heat treatment such as HIP.In this way, strongly textured a ' martensitic microstructure is transformed to equiaxed a grain.Weaker microstructure which leads to get lower thermal stress, also higher elongation break value causes to extend fatigue life [2].Yan et al. [102] stated that the equilibrium aþb phase occurs, martensitic a 0 phase is eliminated and prior b phase is not observed at 1050 C which is above b-transus temperature (when 995 C b-transus temperature considered) for L-PBF Ti6Al4V in the furnace during 4 h.Apparent b regions and dominated a laths instead of equilibrium phase are determined at 950 which is below the b-transus temperature.The same reference article also reported that the compressive strength of Ti6Al4V decreases with above b-transus temperature (see Table 3).On the other hand, heat treatment has the potential not to increase L-PBF cp-Ti strength because of its single-phase (a) [12,68].The comparison of as-built and heat-treated cp-Ti (Grade 2) tubular specimens to measure their endurance limit under the fatigue test showed that the value of endurance limit of heattreated samples became worse at high fatigue strength according to the as-built specimens [31].Zhang and Chen [3] indicated that recent researches are becoming distant from the presence of a 0 phase in L-PBF manufactured Ti6Al4V alloys due to the weak corrosion resistance results despite the heat treatment process.Corrosion resistance is not affected by heat treatment.[119] j o u r n a l o f m a t e r i a l s r e s e a r c h a n d t e c h n o l o g y 2 0 2 2 ; 1 7 : 1 4 0 8 e1 4 2 6 4.2.

Porosity
Looking at the titanium bone implants manufactured via conventional techniques, inadvertent porosity leads to low performance for bone repair and regeneration [120].The fabrications of hierarchical lattice structured bone implants are therefore essential in triggering greater bone regeneration potential [121].The hierarchical lattice structures determines from macro (complex shape, porosity) to micro scale (pore size, interconnectivity) of bone implants [122].High percent of porosity increase the surface area, and this tremendously contributes to the ion homeostasis of the bone.The lattice structure sizes smaller than 1 mm are responsible for bioactivities including protein interactions [123].Highly porous structures allow the movement and the growth of bone cells, as well as the delivery of mineral and oxygen required for bone tissue formation [5] as nutrient and waste transfer (permeability and diffusion) [124].However, a reduction in mechanical strength is observed with larger pore size whereas stress shielding effect is prevented.The porous structure of samples can be designed to match the targeted part properties such as density and mechanical properties (e.g., elastic modulus, hardness, strength) for achieving better compatibility to the bone while facilitating the osseointegration relying on a well-suited porosity level.A pore size between 100 and 700 mm with a 60% of porosity ratio could be   appropriate for biomechanical properties [119,125].Various lattice structures might influence the mobility of bone cells, which can serve as a system to control cell behavior in the implant [126].The most common lattice geometries are surfaces (boundaries) and voxels (discrete volumes) using for porous structured implants [127], some samples of these geometries are the cube, cuboctahedron, tetrahedron [128].However, surface-voxel structures might cause to surface crack initiation at the internal architecture of the porous implant.The self-supporting mathematical surfaces that are named as triply periodic minimal surfaces (TPMS) are focused and researched to solve crack problem, also these surfaces provide precious mechanical and biological advantages because of similarity to trabecular bone and having higher tensile strength at low density.TPMS is an infinite structure and three independent directions with an average zero curvature of the surface (Concave and convex curves are symmetrical at all points) [129].The most known TPMS structures are gyroid, Schwarz diamond, primitive, I-WP which are to form porous complex geometries in FEA analysis [6,19,122].Internal architectures of L-PBF fabricated Ti6Al4V sample as gyroid lattice is shown at Fig. 11.Apart from the TPMS, trabecular bone cell type is another lattice structure which is researched.Ge et al. [118] stated that gyroid and trabecular bone cell types are designed to examine mechanical strength of porous L-PBF manufactured Ti6Al4V.Compressive strength and elastic modulus values of trabecular bone cell type were obtained higher level than gyroid.
As the porosity (volume fraction) rate increases, stress levels differ at the whole part.According to FEA analysis, the centre of unit cells which have the maximum stress levels indicated the higher fracture potential because of bending deformation and tensile stresses [117].Apart from the stress levels, also L-PBF parameter optimization is important for the thin porous parts to prevent uncertainities between CAD model and manufactured part and to reach designed dimensions [23].

Economic and industrial factors
Biomedics is one of the world's rapidly developing industries.It consists of a huge part of the countries' economy and has serious potential funding market.Comparing with AM technology using areas, the medical industry is quite convenient due to integration of AM technologies to tissue engineering and orthopedic applications [7].L-PBF technology is applied to manufacture titanium because of high-level customization, the reduced production time and low material use as cheaper operational cost compared with conventional production techniques.It has been expected that the AM need increases at an annual rate as 21.8% between 2021 and 2028 [130].
Although the manufacturing advantages of L-PBF provides to reduce healing period and implant rejection risk, Velu et al. [7] indicated that the optimization parameter process in L-PBF is still needed to get higher performance of biomedical implants (see Table 4).Also, the gas atomised spherical metal powders, which are used in PBF are mainly expensive according to metal forms for the conventional techniques.To solve this problem, powder product range as metals should be increased and improved, thus the cost is lowered.For instance, more affordable hydrogenerated-dehydrogenerated [108] and hydride-dehydride [106] cp-Ti powders as alternative works to gas atomised spherical powder were developed to produce bone implants via L-PBF (see Table 2).The outcomes of both showed that mechanical strengths and biocompatibility are improved.
Bertoli et al. [42] stated that there are various parameters (laser or electron beam power, layer thickness, gas flow, material quality) included in PBF that affect the final product, however the parameter optimization process makes it more challenging by means of timely and costly manufacturing.To understand the manufacturing process and organize the process parameters under a certain range is also utilized FEA in the simulation of PBF [131].Nevertheless, AM technology supports the improvement of new biomedical implant manufacturing techniques considering with the predictability and repeatability to replace current L-PBF and E-PBF technologies [7].

Conclusions
L-PBF process applications are investigated for cp-Ti and Ti6Al4V biomedical implants.Crucial points concerning the phase transformations and the cooling rates are defined for both materials.The effects of microstructures and surface characteristics of L-PBF fabrication on mechanical, corrosion and biological properties are discussed and compared with currently used conventional cast techniques.Fully densed titaniums with manufacturing L-PBF parameters can improve the biomechanical properties.The defect porosity and the crack propagation however still reveal the need for VED formula optimization in L-PBF process due to the complex melt pool physic.Post-processing methods can give successful results for fatigue resistance, prevention of crack initiations and biological activity for cp-Ti and Ti6Al4V.The evolution of the challenges and opportunities of cp-Ti and Ti6Al4V production using L-PBF technology would enhance functional biomedical implant production limits.Thus, L-PBF might present a broader area for tailor-made titanium bone implant applications not produced with conventional cast techniques.Hierarchical volume fractions (porosity) can be integrated into tailor-made bone implants to provide superior osteointegration and adequate biomechanical properties for cortical and cancellous bones.Obtaining volume fraction design parameters also can be supported by FEA to reduce optimization time and material waste.Integration of novel or potential lattice structures to hierarchical porous implant designs and improvement of manufacturability via L-PBF might enhance biomedical implant coating quality while increasing bone regeneration and reducing the risk of inflammation.
The setting of the L-PBF optimum parameters is still required to eliminate the disadvantages of L-PBF manufacturing of cp-Ti and Ti6Al4V since L-PBF increases the residual stress and the need for heat treatment by means of production methodology.

Fig. 2 e
Fig. 2 e Laser powder bed fusion (L-PBF) Process and Equipment.

Fig. 1 e
Fig. 1 e The cycle of hospital&research center and industry to obtain 3D designed titanium orthopedic implant manufactured with Laser Powder Bed Fusion (L-PBF) technology.

20 Fig. 4 e
Fig. 4 e (Left) Porous Ti6Al4V intervertebral fusion cage placed into resin spine model by Chen [24] is licensed under CC BY 4.0.(Right Bottom) Solid titanium osteosynthesis mini plates by Edelmann [55] is licensed under CC BY 4.0.(Right Upper) Ti6Al4V porous femoral stem for hip replacement by Mehboob [56] is licensed under CC BY 4.0.

Fig. 6 e
Fig.6e Wear rate of cp-Ti with Laser Powder Bed Fusion (L-PBF) and Conventional Cast manufacturing methods[30] (Reproduced with Permission).

[ 106 ]
Minimal impurity was provided under the favor of jet milling and TiO, TiO 2 and Ti 2 O 3 structures formed together in the HDH surface.When HDH cp-Ti powder compared with the high cost atomized powders, similar tensile properties were determined for modified HDH powders and the atomized powders.investigated with cp-Ti scaffolds.Martensite phase via L-PBF.[107](continued on next page) j o u r n a l o f m a t e r i a l s r e s e a r c h a n d t e c h n o l o g y 2 0 2 2 ; 1 7 : 1 4 0 8 e1 4 2 6

Fig. 7 e
Fig. 7 e Physiological environment of body.

Fig. 8 eFig. 9 e
Fig. 8 e Results of histologic assessment additively manufactured cp-Ti and Ti6Al4V porous implants in-vivo femur experiments after eleven weeks.(A) Porous cp-Ti and (B) Zoomed image of cp-Ti.(C) Porous Ti6Al4V and (D) Zoomed image of Ti6Al4V by Wysocki [99] is licensed under CC BY 4.0.

Fig. 10 e
Fig. 10 e Image of balling effect on the surface SLM manufactured commercial pure titanium (cp-Ti) via SEM [71] (Reproduced with Permission).

Fig. 11 e
Fig. 11 e The image of L-PBF manufactured Ti6Al4V part with gyroid unit cells under the optical microscope.
j o u r n a l o f m a t e r i a l s r e s e a r c h a n d t e c h n o l o g y 2 0 2 2 ; 1 7 : 1 4 0 8 e1 4 2 6

Table 1 e
Material properties of cp-Ti and Ti6Al4V for surgical applications.
j o u r n a l o f m a t e r i a l s r e s e a r c h a n d t e c h n o l o g y 2 0 2 2 ; 1 7 : 1 4 0 8 e1 4 2 6

Table 2 e
L-BPF commercial pure titanium parameters and key findings.

Table 3 e
L-PBF parameters and key findings (microstructure, corrosion resistance, mechanical and biological behaviors) of manufactured Ti6Al4V.

Table 4 e
Advantages and disadvantages of L-PBF manufactured TiAl4V and cp-Ti.