Laser and Electron-Beam Powder-Bed Additive Manufacturing of Metallic Implants: A Review on Processes, Materials and Designs Swee Leong Sing,1,2 Jia An,2 Wai Yee Yeong,1,2 Florencia Edith Wiria1,3 1 SIMTech-NTU Joint Laboratory (3D Additive Manufacturing), Nanyang Technological University, HW3-01-01, 65A Nanyang Drive, Singapore 637333, 2Singapore Centre for 3D Printing, School of Mechanical & Aerospace Engineering, Nanyang Technological University, HW1-01-05, 2A Nanyang Link, Singapore 637372, 3Singapore Institute of Manufacturing Technology, 71 Nanyang Drive, Singapore 638075
Received 21 April 2015; accepted 16 October 2015 Published online 29 October 2015 in Wiley Online Library (wileyonlinelibrary.com). DOI 10.1002/jor.23075
ABSTRACT: Additive manufacturing (AM), also commonly known as 3D printing, allows the direct fabrication of functional parts with complex shapes from digital models. In this review, the current progress of two AM processes suitable for metallic orthopaedic implant applications, namely selective laser melting (SLM) and electron beam melting (EBM) are presented. Several critical design factors such as the need for data acquisition for patient-specific design, design dependent porosity for osteo-inductive implants, surface topology of the implants and design for reduction of stress-shielding in implants are discussed. Additive manufactured biomaterials such as 316L stainless steel, titanium-6aluminium-4vanadium (Ti6Al4V) and cobalt-chromium (CoCr) are highlighted. Limitations and future potential of such technologies are also explored. ß 2015 Orthopaedic Research Society. Published by Wiley Periodicals, Inc. J Orthop Res 34:369–385, 2016. Keywords: additive manufacturing; 3D printing; rapid prototyping; selective laser melting; electron beam melting
Additive manufacturing (AM), also commonly known as 3D printing, is a group of processes that join materials to make objects from three-dimensional (3D) model data, usually layer-by-layer, as opposed to subtractive manufacturing methodologies.1 A schematic process chain of a typical powder bed fusion AM process for manufacturing of orthopaedic implants is shown in Figure 1. In recent years, there have been numerous studies applying AM techniques in tissue engineering.2–5 Tissue scaffolds for tissue engineering of bone and cardiac tissues have been fabricated successfully.6,7 With the advancement in AM techniques and materials, the functionality of AM has been extended to the field of orthopaedic implants. One of the key advantages of AM technologies is that they do not have the design constraints that conventional manufacturing techniques have, allowing them to build complex geometries without significant increase in building time. In addition, they require no tooling or moulds and enable the fabrications of several patients implants in the same batch, they are able to provide greater freedom of design to product developers and significantly lower the customization cost.1 According to ASTM F2792-12a, both selective laser melting (SLM) and electron beam melting (EBM) are classified as powder bed fusion technologies where thermal energy selectively fuse regions of a powder bed.8 In this paper, the application of these two AM technologies in fabrication of orthopaedic implants is reviewed, with specific focus on biocompatible materials such as 316L stainless steel, titanium-6aluminium4vanadium (Ti6Al4V) and cobalt-chromium (CoCr).
Correspondence to: Wai Y. Yeong (T: þ65 6790 4343; F: þ65 6791 9690; E-mail:
[email protected]) # 2015 Orthopaedic Research Society. Published by Wiley Periodicals, Inc.
ADDITIVE MANUFACTURING OF ORTHOPAEDIC IMPLANTS Powder bed fusion processes use an energy source to melt and fuse selective regions of powder according to computer aided design (CAD) data. When the selective melting of one layer is completed, the building platform is lowered by a predetermined distance (usually 20–100 mm for SLM and 100 mm for EBM) and a next layer of powder is deposited on the platform. The process is then repeated with successive layers of powder until the required part is completely built.1 The exclusion of sacrificial binders in the process enables near-full density parts to be built. This gives it an important advantage over binder-based processes in direct part manufacturing. An overview of key steps within powder bed fusion AM is shown in Figure 2. Two of the more representative powder bed fusion techniques are SLM and EBM. Selective Laser Melting Selective laser melting (SLM) uses a fibre laser source as the energy source.9 The whole process is carried out in an inert gas filled chamber which ensures higher purity by minimizing the oxygen in the environment and reduces the risk of hydrogen pick up. The schematic of the SLM system is shown in Figure 3. The SLM system comprises of an fibre laser, which can operate up to 1 kW depending on the laser module installed in the system.1 The beam focus is controlled by the galvanometer and the movement of the beam on the build table is controlled by F-theta lens. For building a part, a powder layer of 20–100 mm thickness is spread over the table. The powder is carried and spread by the powder recoater across the build table. The build table can be preheated up to 200˚C. The selective melting of the powder layer based on the geometry defined by the CAD file is done by the laser. In SLM, every layer of a part is built in two steps. The outer boundary of the part is built first which is JOURNAL OF ORTHOPAEDIC RESEARCH MARCH 2016
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Figure 1. Process chain for powder bed fusion AM for orthopedic implants. The pre-processing before AM includes 3D modelling, file repairing and preparation, and lastly, slicing of the 3D CAD into layers. Post-processing may include shot peening, heat treatment and polishing of fabricated parts.
referred as contouring and the powder within the contour is melted subsequently to complete one layer. This process continues until the desired threedimensional part is fully completed.10 SLM has been utilized to fabricate orthopaedic implants such as replacements for zygomatic bone11 and finger.12
Electron Beam Melting Electron beam melting (EBM) is another metal additive manufacturing technique which is believed to revolutionize the implant manufacturing industry. This process is conceptualized and patented by Arcam AB based in Sweden. The process utilizes electron beam energy to melt the metal powder.13 Since
Figure 2. Key steps in powder bed fusion AM. The key steps of the processes includes powder deposition, either by a recoater or powder rake, followed by irradiation using the energy source and lastly, lowering of the build platform. The cycle then repeats until the part is fully built. JOURNAL OF ORTHOPAEDIC RESEARCH MARCH 2016
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Figure 3. Schematic of SLM system. The key components of SLM include the laser system (a fiber laser, F-theta and galvanometer used to control the laser beam movement) and the mechanical system (movable build platform and powder recoater). A view of laser scanning of the powder bed is also shown.
electron beam is used as the energy source the entire process takes place in a vacuum chamber. The vacuum ensures high purity by providing oxygen free environment and reduces the risk of hydrogen pick up. This feature is extremely beneficial in fabricating parts out of titanium-6aluminium-4vanadium (Ti6Al4V) because the low levels of interstitial elements can be controlled during fabrication. Also, while building the part, an elevated temperature of about 700˚C is maintained in the chamber to reduce residual stresses and hence, distortion and warpage. A schematic of EBM system is shown in Figure 4. The EBM system comprises of an electron gun, similar to an electron gun in a scanning electron microscope or an electron beam welding machine, operates at a power of 60 kW to generate a focused beam of energy density above 100 kW/cm2. The beam focus is controlled by the electromagnetic lenses and the movement of the beam on the build table is controlled by deflection coils. For building a part, a powder layer of 100 m thickness is spread over the table. The powder is supplied from two hoppers kept inside the build chamber. A moving rake fetches powder from both sides and spread over the table. The electron beam first pre-heats the powder layer with a higher scan speed, followed by melting the powder layer based on the geometry defined by the CAD file. In EBM, every layer of a part is built in two steps. The outer boundary of the part is built first which is referred as contouring and the powder within the contour is melted subsequently to complete one layer.
This process continues until the desired three-dimensional part is fully completed.9 EBM has been used to produce orthopaedic components such as knee, hip, jaw replacements, and maxillofacial plates.14–17 EBM produced implants, such as acetabular cups, have also gained approval from United States Food and Drug Administration (FDA) and are CE-certified since 2010 and 2007, respectively. These two processes share similar advantages and are actively considered for manufacturing of orthopaedic implants. These advantages include: Fabrication of complex products, novel shapes, hollow structures, and functionally gradient materials that are not otherwise feasible. Efficient approach that reduces production costs and speed time-to-market for high-value components. Excellent material properties, almost no porosity, possibility to combine different materials, significantly reduced material waste. Elimination of expensive tooling. Due to these advantages, numerous studies have been carried out by different researchers to prove the suitability of these processes in producing implants as shown in Figure 5. Orthopaedic implants are mostly used for structural reinforcement, as an artificial bone, which is inserted inside the corpus. They include both temporary implants, JOURNAL OF ORTHOPAEDIC RESEARCH MARCH 2016
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Figure 4. Schematic of EBM system. The system includes an electron beam system (electron gun assembly, electron beam focusing lens and deflection coils used to control the electron beam) and the mechanical system (movable powder rake and fixed powder cassettes). A view of electron beam scanning the powder bed is also shown.
such as plates and screws, and permanent implants that are used to replace body parts such as hip, knee and fingers directly.18 Permanent implants put more emphasis on strength, toughness, abrasion resistance in artificial joints as well as tribology.
DESIGN CONSIDERATIONS FOR ORTHOPAEDIC IMPLANTS USING ADDITIVE MANUFACTURING While exploiting the advantages of SLM and EBM in fabricating functional permanent implant, there are various design constraints that should be taken into consideration for designing an optimal implant. In the subsequent sections, several critical design factors such as the need of data acquisition for patient-specific design, design dependent porosity for osteo-inductive implants, surface topology of additive manufactured
implants and design for reduction of stress-shielding in implants are discussed. Data Acquisition for Patient-Specific Design One of the main advantages of using AM for manufacturing implants is the ability to customize the implants according to patient’s specifications quickly. The manufacturing of custom made implants necessitates the recording of the patient’s anatomical data by scanning processes and medical imaging such as magnetic resonance imaging (MRI) and radiography (X-ray). The patient’s anatomical data is then reconstructed in 3D through medical image processing and implants geometry modelling such as addition of features that are specific and achievable by AM. The data handling process requires specific skillset and
Figure 5. Samples of (a) hip joint by SLM after polishing and (b) as-fabricated acetabular cup by EBM fabricated in-house, with magnified view of the porous lattice structure. JOURNAL OF ORTHOPAEDIC RESEARCH MARCH 2016
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software applications. A schematic of the process flow in data handling prior to additive manufacturing of implants is shown in Figure 6. Mazzoli et al. fabricated cranial plates in titanium for repair of the skulls.15 The anatomy of the defect was reproduced using commercial medical modelling software which interfaces computer tomography (CT) and CAD data. Based on the 3D visual models, implant design was virtually carried out using haptic freeform modelling which involves the direct interaction between user and virtual model. Prior to manufacturing the final implant, the design of the implant was validated for quality of fit, shape and symmetry with replicas manufactured by selective laser sintering (SLS). A coordinate measuring machine was used for validation to ensure model accuracy. Then the implants were directly manufactured using EBM for implantation. In another case, Jardini et al. fabricated a customized Ti6Al4V implant for surgical reconstruction of a large cranial defect.16 Similar to the previous case, a series of computed CT data was obtained for extraction of the cranial geometry. The prosthetic implant was then fabricated using SLM. During the surgical procedure, which took 3 h, the implant was then placed into the patient. When non-customized implants are used in similar surgery, the duration is approximately 6 h. The reduction in duration of surgery is due to preoperative planning of correct geometrical and anatomical details.16,17 The case studies above showed the need of efficient and validated data acquisition in order to reconstruct the orthopaedic implants with accurate fit via SLM and EBM.
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Design Dependent Porosity for Osteo-Inductive Implants Bone is a composite material in nature which contains about 45–60% minerals, 20–30% matrix, and 10–20% water.19 The human skeleton consists of two types of bone, namely, cortical, and trabecular bones. These two types of bone differ in terms of proportions of organic and inorganic materials, degree of porosity and organization.19 Due to these differences, orthopaedic implants require precise designs of pores and porosities in order to mimic the bone properties closely.2 A porous implant material with adequate pore structure and appropriate mechanical properties has long been sought as the ideal bone substitute as interconnected pores allow tissue in-growth and thus anchor the implant to the surrounding bone, preventing loosening of the implant.20 Recent publications have shown that micro-porosity is an essential element in osteo-induction by biomaterials. The pores of the structure also have to be interconnected in order to ensure the bone in-growth and cell migration.21–23 Open porosity is essential for osteo-induction (promotion of the differentiation of progenitor cells down an osteo-blastic linage), osteo-conduction (support bone growth and encourage the in-growth of surrounding bone) and osteo-integration (integration into surrounding bone). In addition, the porosity of the implants and the pore architecture are critical for encouragement of cell migration.3 Currently, components for orthopaedic use have an average pore size ranging from 400 to 600 mm with a volume porosity of 75–85%.24 In addition to having suitable open pore size and porosity for cell attachment and proliferation, the implant must have sufficient mechanical strength to support physio-
Figure 6. Process flow in data acquisition for custom made implants by AM. This includes scanning and medical imaging by X-ray or MRI, followed by reconstruction of raw scanning data into 3D model. With the 3D model, AM specific designed can be incorporated into the CAD, and the final file can be prepared. Data acquisition is then followed by the pre-processing in AM using the finalized CAD model. JOURNAL OF ORTHOPAEDIC RESEARCH MARCH 2016
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logical loadings in vivo.2,23 Various researches have been carried out in areas relating porous structures fabricated and SLM or EBM and their biological responses, the key results are summarized in Table 1. In addition to pore size and porosity, recent studies indicated that the shape of pores may also affect cell proliferation and differentiation.27,32 Van Bael et al. investigated the effect of three different pore shapes. The pore shapes affect the permeability of the cells and a higher permeability implies that the cell suspension has less resistance when permeating the scaffold,
leading to lower time for cell attachments. Amount of pore occlusion is higher on hexagonal pores compared to triangular or rectangular pores. However, bridging behavior of the cells is not dependent on pore shapes.27 Surface Topology and Coating of Additive Manufactured Implants In this section, important biological responses of AM implants based on surface roughness and grain size are discussed. Biological responses such as osteoblastic cell adhesion, growth and differentiation are
Table 1. Biological Responses of Porous Structures Fabricated by SLM and EBM Process Material
Pore Size (mm)
Porosity (%)
SLM
Ti6Al4V
240–730
68 and 88
SLM
Ti6Al4V
450–1200
SLM
Ti6Al4V 500 and 1000
SLM
Ti6Al4V
280–420
EBM
Ti6Al4V
500–1500
60–83
EBM
Ti6Al4V
500–700
65–70
EBM
Ti6Al4V 640 and 1200
Post Processing
Soaked in HF for 10 min, rinsed with demineralized water and ethanol. Surface coated with octacalcium phosphate or hydroxyapatite, both doped with Mg
73 and 82
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Soaked in HNO3 for 30 min, followed by washing in tap water and sonication in 50/ 50 mixture of ethanol and deionized water Blasted with Ti6Al4V powder
Powder blasting, rinsed ultrasonically with acetone, alcohol and distilled water for 30 min. Sterilized with steam autoclaving for 30 min
Key Findings
Reference
Scaffolds provided enough mechanical support and encouraged bone formation. Porous structures allow total overgrowth of all pores of 500 mm and a significant proportion in the range of 500 to 600 mm by osteoblasts. Smaller pore size results in higher cell attachments due to lower permeability.
25
Both as build and surface modified implants are biocompatible and integrate with the bone with good boneimplant bonding. Volume of bone was highest for implants coated with hydroxyapatite. Collagenous extracellular matrix can be deposited on the surface and within the pores of structures.
28
High degree of osseo-integration was observed inside the porous structures after 26 weeks. Structures promote soft tissue healing. Structures supported the attachment and proliferation with minimal inflammatory cytokines secretion. Smaller pore size were more compatible and better facilitate osteo-genesis.
30
26
27
29
31
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related to surface energy and roughness,33 while grain size of the material will affect the corrosion performance of the implant. The degree of osteo-integration relies on material dependent factors such as surface topography. The as-fabricated SLM or EBM part can have a rough surface with a Ra value of several microns, which is deemed to be more associated with bone-implant contact area and interface strength. Surface appearance was found to be influenced by the powder size.34 Traditionally, the surface roughness value of an implant can be modified by sandblasting, acid etching or plasma spraying. Biemond et al. used EBM to produce implants with different surface topographies and the friction coefficients were in-vitro compared with plasma sprayed titanium implants and sand blasted implants. The frictional and bone in-growth properties of EBM produced surface structure are found to be similar to the plasma-sprayed surface. A comparative study on bone in-growth potential of both electron beam and laser beam fabricated trabecular bone implant has been studied by Biemond et al.35 The original design of the trabecular implant surface was similar for both production techniques. However, the SLM and EBM specimens that resulted after production were different in gross morphology and surface texture because of the differences in processing conditions. Based on all histological and push-out test results they found that there are no large differences in in-growth potential between the trabecular bone implants made by the two processes. Li et al. evaluated the biological properties of EBM produced Ti6Al4V implants with biomimetic coatings in vitro and in vivo. The implants were produced with open pore size ranging between 500 to 600 mm. The in vitro culture of osteoblasts on the porous Ti6Al4V implants has shown a favourable circumstance for cell attachment and proliferation as well as cell morphology and spreading. In vivo histological analysis has obtained a rapid in-growth of bone tissue from calvarial margins toward the center of bone defect in 12 weeks. They found that the bone in-growth within the implant with or without apatite coating was quite similar, which means the implants can be directly used without surface modifications.36 De Wild et al. suggested that the as-fabricated SLM or EBM part should be surface treated before use, because bone formation and bridging had been found significantly increased in sand blasted and acid-etched implants compared to the untreated.37 Pattanayak et al. also reported that bone affinity of chemical and heat treated porous bodies (smoother surface) were significantly higher than that of untreated implants.38 However, recently, there is a school of thoughts that SLM or EBM fabricated parts may be directly used without surface treatment. Thomsen et al. compared EBM-made Ti6Al4V implants with conventionally wrought machined implants in vivo and found no appreciable difference in tissue response in short term (6 weeks) between implants made by these two methods.39
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Biemond et al. found that the bone-implant contact of EBM produced implant is around 24–25%, similar to titanium plasma sprayed implant (25.5%),40 and that the presence of coating did not produce better bone in-growth or mechanical fixation strength, 35 which suggests that the as-fabricated EBM part may be directly used without surface treatment. Bertollo et al. also found no difference in bone in-growth properties by comparing asfabricated EBM parts with titanium plasmasprayed ones.41 In a study, porous Ti structures were fabricated using SLM and the surfaces were left untreated, sandblasted, or sandblasted/acid etched. It was observed in the rabbit model that bone bridging was significantly increased in sand blasted acid-etched scaffolds, indicating the needs for specially designed surface characteristic to improve osteo-conduction of the implants.37 The contradicting conclusion from various studies may be explained by the surface roughness of the implants studied, as shown by Ponader et al.42 The study investigated the responses of osteoblasts to the surface of EBM implants and they found that when Ra is below 24.9 mm, surface roughness has a positive effect and can promote the proliferation and differentiation of human osteoblasts, but when Ra exceeds 56.9 mm, surface roughness has a negative effect on the proliferation.42 This might offer an explanation for the aforementioned puzzling findings as well as a suggestion that when surface treatment should be considered for SLM and EBM parts. If the as-fabricated implant already has a Ra value below 24.9 mm, the surface treatment may be unnecessary. Xin et al. found the corrosion behaviour of SLMmade CoCr dental implants is pH dependent. At pH 5, no differences were observed between SLM implants and traditional cast implants, but at pH 2.5, SLM implants had a better anti-corrosion behaviour.43,44 The different corrosion resistances of SLM and cast samples under more acidic conditions may be due to difference microstructures. The cast part has microstructure of a heterogeneous two phase mixture consisting of a solid solution and a crystalline phase. On the other hand, the SLM part showed a compact homogeneous cellular microstructure with a much smaller gain size than cast part. The smaller grain size accounts for the better corrosion resistance in acidic conditions.44 These studies conclude that with careful control of the mechanical and surface properties, SLM and EBM offer the potential to produce implants with unmatched properties and biological responses, on par or more superior than conventional implants. Design for Reduction of Stress-Shielding In order to address the issue of stress shielding, many design and manufacturing techniques have been tried by considering flexibility and interface properties. Some of the designs include hollow stems, grooved JOURNAL OF ORTHOPAEDIC RESEARCH MARCH 2016
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stem and modular stems. Harryson et al. carried out design and fabrication of titanium hip implants using EBM to reduce the stress shielding.45 The aim was to reduce the implant’s bending stiffness while maintaining its mechanical strength. They fabricated hip stems with mesh configuration, hole configuration and as complete solid. The stress distribution on the femurs was found to be different for all the three cases. The proximal portion of the femur exhibited higher stresses with decreasing stem stiffness. Among the three configurations, Ti6Al4V stem with mesh results in a more even stress distribution in the proximal portion of the femur. The more even stress distribution most likely leads to a reduction in bone remodelling and hence, the lowest level in stress shielding.45 Therefore, in SLM and EBM, the newly created bulk structure can be represented by a range of porosity and pore size, which can be tailored to reduce the stiffness of a metal implant to overcome the problem of stress shielding.13,46–48
ADDITIVE MANUFACTURED MATERIALS FOR ORTHOPAEDIC IMPLANTS Some of the common metals used in orthopaedic applications include 316L stainless steel, Ti6Al4V and CoCr alloy.23 316L Stainless Steel This material is low-cost and easily available, making it a suitable in the medical industry as a biocompatible metal bone implant. Together with AM, it is well suited for these applications as implants or prostheses can be individualized with very low customization costs. Yang et al. investigated the optimization of building accuracy and density of orthodontic products using a self-developed SLM machine and they were able to achieve the required surface quality and mechanical properties.49 Li et al. studied the possibility of making SLM 316L stainless steel parts with gradient porosity where the dense portion is designed for strength and the porous part is designed to enhance tissue growth in biocompatible implants.50 Bibb et al. reported SLM denture frameworks with the same material.51 Bibb et al. also presented 4 case studies on surgical guides in different maxillofacial (jaw and face) surgeries.52 Kruth et al. published on a biocompatible metal framework for dental prostheses53 and Wehmoller et al. reported body implants of cortical bone, mandib-
ular canal segment and support structures or tubular bone made from SLM 316L stainless steel.54 The processing parameters for fabricating parts using 316L stainless steel in EBM are currently under development. Details of research in this area are not available in open literature. Mechanical Properties Relative Density The relative density of a part is often used as an indicator of the quality of the SLM produced part. Relative density is the ratio of the density of the SLM fabricated part to the theoretical density of the bulk material. Tolosa et al. was able to achieve a relative density of 99.90%55 and Yasa et al. achieved 99.95% relative density with laser re-melting.56 It is of interest to note that laser re-melting increases the power consumption as well as time for the fabrication process as each layer is scanned twice. Strength and Hardness Steel is often used in many applications due to its strength. It is reported that SLM produced steel components are stronger and less malleable compared to the forged counterparts. The rapid cooling in SLM process results in refined microstructure which enhances the tensile strength with a reduction in ductility. The ultimate tensile strength, yield strength, elongation and micro-hardness of 316L stainless steel parts produced by SLM and forging are shown in Table 2. A fine cellular dendritic structure is characteristic of SLM produced 316L parts, due to the rapid solidification during the SLM process. This results in the higher strength obtained as compared to forged 316L stainless steel parts which experience lower cooling rates. Surface Roughness It is common to achieve roughness of about 20 mm via SLM. Post processing such as sand-blasting, shotpeening or manual grinding is often needed to achieve a smooth and shiny surface. However, for 316L stainless steel, Delgado et al. were able to achieve surface roughness of 5.82 mm without any post processing.63 Kruth’s group, who has been experimenting with laser re-melting, was able to achieve surface roughness as low as 2 mm for 316L stainless steel.56 Titanium-6Aluminium-4Vanadium (Ti6Al4V) Most research on titanium and its alloys is driven by its potential application as body prostheses due to
Table 2. Tensile Properties and Micro-Hardness of 316L Stainless Steel by SLM and Forging Properties Ultimate tensile strength (MPa) Yield strength (MPa) Elongation (%) Micro-hardness (HV) JOURNAL OF ORTHOPAEDIC RESEARCH MARCH 2016
SLM
Reference
Forging
Reference
480–760 350–640 10–30 220–279
53,57,58
450–818 150–230 50–62 133–140
59,60
57,61 53,61 53,58
59,62 59,60 59
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their biocompatibility. Ti6Al4V are of high interest because of their applications in both aerospace and biomedical industries.64,65 This group of metallic materials has been widely used for various orthopaedic implants due to their good biocompatibility, superior corrosion resistance and high mechanical strength.38,66 Moreover, they have high specific strength and elastic moduli closer to bone than CoCr alloys and stainless steel.67 Performance requirements for implants made in Ti6Al4V alloy are specified by ASTM standards and US FDA. Ti6Al4V components can be produced with variety of microstructures depending on the method used for processing the alloy. For example, casting, wrought ingots and powder metallurgy give three different microstructures for Ti6Al4V. This is because for pure Ti, the microstructure is completely a. When pure Ti is alloyed with a and b stabilizers, b phase forms along the grain boundary. The percentage of a and b phases varies depending upon the processing conditions such as the temperature, cooling rates and degree of mechanical working. Research into Ti6Al4V body implants using SLM have been done by several groups. Lin et al. studied the structure and mechanical properties of a Ti6Al4V cellular inter-body fusion cage,68 Murr et al. focused their attention on the microstructure and mechanical properties of SLM Ti6Al4V for biomedical applications69 and Warnke et al. conducted cell experiments and showed that SLM Ti6Al4V porous scaffolds allow total overgrowth of osteoblasts (bone cells).26 Vandenbroucke & Kruth examined the dimensional accuracy of the SLM process for fabrication of dental frameworks.70 In recent years, various research groups have also carried out numerous studies on SLM of Ti6Al4V for body implants.22,25,65,71–82 In one of the latest studies, Biemond et al. examined the bone in-growth potential of trabecular-like implant surfaces produced by SLM of Ti6Al4V in goats and concluded that the SLM produced parts showed good bone in-growth characteristics after 15 weeks.35 Ti6Al4V is also well studied for EBM. Detailed study on microstructural and mechanical characteristics of Ti6Al4V parts made using EBM have been done by many researchers.32,35,41,83,84 All the studies reported that the microstructure of EBM produced Ti6Al4V parts consist of a phase with small amount of retained b phase. This is different from the microstructure obtained by SLM, which is completely martensitic. This microstructural difference is attributed
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to the difference in processing conditions. Consequently, the mechanical properties also differ slightly for Ti6Al4V parts produced in EBM and SLM. Mechanical Properties Relative Density SLM and EBM of Ti6Al4V have been very positive in terms of the relative density attained by various researchers. The highest relative density reported for SLM Ti6Al4V is 99.80%.70 Arcam AB reportedly obtained fully dense parts for Ti6Al4V by EBM. Strength and Hardness Table 3 shows the highest reported ultimate tensile strength for Ti6Al4V, along with their respective yield strength and elongation for SLM, EBM and casting produced parts. The tensile strength and micro-hardness of EBM produced Ti6Al4V parts are lower as compared to SLM produced Ti6Al4V. This is due to the difference in the microstructure of the parts produced by the two processes. Microstructural evolution is primarily a function of cooling rate and materials processed in SLM and EBM undergo different cooling rates. Very high cooling rates observed in SLM resulted in a harder martensitic phase as compared to the softer a phase obtained from EBM, where the cooling rates are lower. In EBM the material cools down to a temperature of between 650 and 700˚C which is the temperature of the build chamber. The vacuum build chamber in EBM also results in lower cooling rate. This does not allow the transformation of a to martensitic phase in EBM.9 In comparison, the build chamber temperature of SLM is about 30–60˚C, which create a steeper thermal gradient, and the flow of inert gas, creating heat transfer by convection in the build chamber, resulting in higher cooling rates. Therefore, the difference between EBM and SLM in processing Ti6Al4V lies in the final microstructure which in turn, affects the mechanical properties. Cobalt-Chromium (CoCr) CoCr has been studied by various groups for implant applications. Oyague et al. and Kim et al. separately evaluated the fit of dental prostheses produced by SLM and reached different conclusions about the suitability of SLM technology in producing dental prostheses.89,90 In terms of hardness, elastic modulus and strength, Ayyildiz et al. concluded that CoCr produced by laser AM is suitable for dental applications.91
Table 3. Tensile Strengths and Micro-Hardness of SLM, EBM, and Casted Ti6Al4V Properties Ultimate tensile strength (MPa) Yield strength (MPa) Elongation (%) Micro-hardness (HV)
SLM
Reference
1250–1267 1110–1125 6–7 479–613
70,85 70,85 70,85 10,71
EBM
Reference
Casting
Reference
830–1150 915–1200 13–25 358–387
84,86
934–1173 862–999 6–7 294–360
87,88
84,86 84,86 84,86
87,88 87,88 89,90
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Table 4. Tensile Strengths and Micro-Hardness of SLM, EBM, and Casted CoCr Properties Ultimate tensile strength (MPa) Yield strength (MPa) Elongation (%) Micro-hardness (HV)
SLM
Reference
562–884 951–1308 10.2–16.4 458.3–482.0
93–95 93–95 93–95 43,91
Mechanical Properties Relative Density CoCr achieved a relative density of 99.94% by SLM.92 Arcam AB reportedly obtained fully dense parts from CoCr by EBM. Strength and Hardness Table 4 shows the highest reported ultimate tensile strength for CoCr, along with their respective yield strength and elongation for SLM and casting produced parts. New Materials Magnesium and its alloys have great potential for orthopaedic applications as they have been shown to be fully bioresorbable, have mechanical properties aligned to bone and induce no inflammatory response. In addition, they are also osteo-conductive, encourage bone growth and have a role in cell attachment.19 Man’s group examined the SLM of magnesium for applications in light-weight biodegradable and bioresorbable magnesium parts for orthopaedic implants.97–99 There are recent studies that aim to replace Ti6Al4V with other titanium based alloys. Chlebus et al. studied the possibility of using Ti-6Al-7Nb for medical implants as it replaces vanadium with niobium in its chemical composition and this alloy has higher corrosion resistance and bio-tolerance compared to Ti6Al4V.100 Further studies on this novel titanium alloy were done by Marcu et al. on endosseous implant.101 Szymczyk et al. also examined cultured cell growth of Staphylocuccus aureus on Ti6Al-7Nb scaffold which have demonstrated the potential of this titanium alloy in this application.102 SLM of Ti-24Nb-4Zr-8Sn has been examined by Zhang et al. as an improvement over Ti6Al4V as it is an alloy of lower modulus. This results in a closer match of moduli between implant and surrounding bone and thus have higher possibility of preventing bone resorption, which causes implant loosening.97 Tantalum is a biomaterial with unique mechanical properties. With designed porosity, it can achieve an elastic modulus similar to that of bone, which minimises stress shielding.19 Thijs et al. experimented using pure tantalum for fabrication using SLM and achieved a relative density of 99.6 %. The yield strength of SLM Ta is also reported to be higher than those obtained via conventional methods.103 Ti6Al4V and CoCr are the materials which are commercialized for application in EBM machine, proJOURNAL OF ORTHOPAEDIC RESEARCH MARCH 2016
EBM (After Heat Treatment)
Reference
Casting
Reference
960 560 20 325.5
Arcam Arcam Arcam Arcam
296–568 296–568 8.0–10.7 324.0–384.8
93,94 93,94 93,94 43,96
cess development is currently undergoing for other materials such as NiTi, 316L stainless steel, titanium aluminides, copper, and tool steels. However, there is an obstacle in expanding the material library for this process at current stage. This is due to EBM not being an open process where the materials that can be processed by the technique are controlled by the original equipment manufacturer, Arcam AB.
CHALLENGES, POTENTIAL AND CURRENT ACTIVE RESEARCH IN ADDITIVELY MANUFACTURED ORTHOPAEDIC IMPLANTS As mentioned previously, for the additively manufactured implants to be suitable for implantation, porosity is necessary. Formation of pores can occur inherently during the SLM and EBM processes. It is known that when interaction of laser or electron beam with metal powder occurs, the energy is absorbed by powder particles through bulk coupling and powder coupling mechanisms. The deposit of high energy leads to a high temperature rise and a large amount of liquid formation. Simultaneously, a steep temperature gradient is formed, leading to surface tension gradient as well as the resultant shear stress and convective movement of the melt pool which is known as Marangoni effect.104 During the subsequent rapid solidification process, gaseous bubbles formed can float to the melt surface due to Marangoni flow and escape from the melt pool or trap in the solidified structure, resulting in pores. However, this uncontrolled formation of pores is undesirable and can be minimised by the careful selection of processing parameters such as laser power, scan speed and hatch spacing. Fortunately, EBM and SLM provide the opportunity for fabrication of cellular lattice structures to achieve the desirable designs and levels of porosity. However, there is a limit to which osteo-inductive potential can be increased by increasing the micro-porosity as a mechanically stable surface of the material is needed in order to facilitate new bone formation.21,68,105,106 Furthermore, the current implants composed of a single material, sometimes with a coating layer, which is essentially uniform in composition and structure in the longitudinal direction.107,108 This leads to constant properties, such as strength and biocompatibility throughout the implant. However, a single composition with uniform structure cannot satisfy all the requirements needed for implants.109 For specific bone tissues, such as long bones, they are structurally organized in
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Figure 7. Porous structure generated by CASTS. The cellular porous structure is generated from repeating unit cells allowing designed porosity across the structure.
such a way that bone porosity varies from the outer perimeter to the inner section in order to achieve normal bone function. In this case, the mechanical strength of the bone decreases gradually from the outer to inner regions and thus can be regarded as a functionally graded structure.2 Therefore, the concept of functionally graded materials (FGM)107 and multimaterial processing in SLM and EBM may be suitable for implants. These parts possess the distinguishing feature of non-homogeneity with regard to strength related properties including yield strength, fracture toughness, fatigue and creep behaviour.110 Challenges in Fabrication of Cellular Lattice Structures Cellular lattice structures can be classified into structures with stochastic and non-stochastic geometries. Stochastic lattice structures have random variations in the shape and size of the cells, whereas, nonstochastic or periodic lattice structures have repeating lattice structures and can be categorized by their shapes and sizes.111 There are many ongoing researches on the fabrication of cellular lattice structures by SLM and EBM, focusing on the dimensional accuracy of the fabricated structures,112–114 mechanical properties48,83,112,115,116 and biocompatibility of these structures.29,37,117 Researches have also been done on development of automated algorithm for cellular lattice structures. One of the in-house developed software, called the “computer-aided system for tissue scaffolds” (CASTS), can assemble 13 designed unit cells 2 into porous structures which can be modified easily.3 The details of CASTS can be found in work done by Chua et al.118 CASTS allows the designs with size and shape adapted to particular patient to be produced by AM. Examples of porous structures designed using CASTS are shown in Figure 7. The examples show that porous structures with controlled porosity can be generated directly using CASTS. Samples of cellular lattice structures fabricated using SLM is shown in Figure 8. The samples shown clearly demonstrated the capability of powder bed fusion AM in producing cellular lattice structures from CAD models in different shapes. However, key challenges have also been identified in these areas, such as the powder adhesions to the struts38 and the difficulty in removal of the unmelted powder within the structures.111
Functionally Graded Materials FGM are composite materials formed from two or more constituent phases with a continuously variable composition or structures.119,120 In addition, FGM has the potential to eliminate the problem of bone remodelling due to mismatch in mechanical properties. FGM signifies a new class of composites, which consists of a graded pattern of material compositions, microstructures or macrostructures, allowing better matching of corresponding mechanical properties.121,122 The resulting FGM can have a microstructure or macrostructure that produces continuous or discrete change in mechanical properties123 due to variation in composition or structure changing gradually.108 The main advantages of using FGM implant are reduction in stress shielding effect on the surrounding bones, improvement of biocompatibility with bone tissues and meeting the biomechanical requirements at each region of the bone while enhancing bone remodelling and maintaining the bone health.121 SLM and EBM permit the creation of very complex geometries with a gradient of porosity perpendicular to the longitudinal axis of the implant,108,124 which in turn allows the choice of property distribution to achieve required functions. A sample of functionally graded porous structure, with increasing porosity from left to right, fabricated using SLM is shown in Figure 9. The fabricated sample has shown the versatility of SLM in producing porous structures and FGM with wide range of porosities. Cheng et al. fabricated cellular structures that have comparable compressive strength similar to those of trabecular and cortical bone, with elastic modulus between 0.2 and 6.3 GPa.83 Watari et al. studied FGM using titanium and hydroxyapatite (HAP).107,109 The samples were sintered using electric furnace heating, high frequency induction heating and spark plasma sintering. The study concluded that gradient functions in both biochemical affinity to osteogenesis and the mechanical properties with stress relaxation effect could contribute to efficient biocompatibility.107 It also demonstrated that tissue reaction changed in response to the changes in gradient composition or structure of materials, which implied that tissue response could be controlled using FGM.107 Lin et al. studied design of dental implants using FGM.121 The study found out that by lowering the FGM gradient, a better performance in bone turnover could be achieved, however, JOURNAL OF ORTHOPAEDIC RESEARCH MARCH 2016
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Sing et al. also achieved good interfacial characteristics between AlSi10Mg aluminium alloy and C1800 copper alloy using the same setup and method.131 Terrazas et al. investigated multi-material material processing in EBM with Ti6Al4V and pure copper. Good metallurgical bonding was also observed between these materials, however, undesirable thermal stresses, due to differences in physical properties, and formation of brittle intermetallic phases upon joining of both materials are observed.132 Nonetheless, these studies open up the opportunity to explore other alloy systems that are suitable for orthopaedic applications, such as the stainless steel-Ti6Al4V and Ti-Ta systems.
CONCLUSION AND OUTLOOK Figure 8. Samples of cellular lattice structures fabricated by SLM. Tubular and disc shaped cellular lattice structures produced showing the capability of generating porous structures in various shapes and sizes.
this would reduce the stiffness of implantation which would lead to higher risk of damage during the early healing stage.121 Multi-Material Processing One of the limitations of the SLM and EBM processes is that a single build is usually restricted to one type of materials. The machines commercially available are designed to work with only one material per build. Even though, for the EBM system, there are two containers for metal powder, located at each side of the build chamber, the system rakes from both directions. This causes mixing of the two powder materials. However, for metallic implants, due to the varying requirement at different parts of the implants, there is often a need for multi-material design in a single structure. Multi-material processing has been carried out by AM systems that use directed energy deposition process such as Optomec’s laser engineered near net shape (LENS) system125–127 and selective laser sintering (SLS) systems.128,129 Unfortunately, there is little prior research done on multi-material processing of metal via SLM or EBM and the properties of such parts are not fully understood.130 Recently, Liu et al. experimented with multi-material processing in SLM using 316L stainless steel and C18400 copper alloy. Good metallurgical bonding was obtained at the interface between these two alloys.58
With the advancement in AM technologies, it is now able to fabricate fully functional parts directly. In particular, SLM and EBM provide the orthopaedic field the opportunities to mass customize implants at a lower cost due to their ability to fabricate parts with complex and intrinsic designs that are specific to individual patients. To summarize, all the above studies show the immense potential of EBM and SLM to become the more preferred method for producing orthopaedic implants in future. However, there are still research challenges that need to be overcome, in areas such as: Material Research Current EBM research are limited to narrow range of materials due to the close system enforced by the original equipment manufacturer, however, if the technology follows the research trend of SLM, the material library of these technologies will become wider and allows for some exciting research in the field in the near future. Biocompatibility of SLM and EBM Produced Orthopaedic Implants Although considerable amount of work has been done on the biocompatibility aspects of EBM and SLM produced implants, detailed studies on the long term performance based on the fracture mechanics aspects is lacking. The duration to failure of implants is another matter of concern. Revision surgery after failure of an implant creates lot of inconvenience and unnecessary pain to the patient. A long lasting implant could be achieved by selecting the optimal combination of surface roughness, pore size and porosity. Furthermore, the customized implants result in
Figure 9. Sample of functionally graded porous structure fabricated using SLM with porosity increasing from left to right. The porosity of the structures can be functionally graded and varied to match the porosity of neighbouring bones after implantation.
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close fitting and reduce operating time, the probability of rectification or revision surgery is also reduced. These can be achieved by the careful selection on processing parameters of SLM and EBM as well as the designs of the implants.
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Designed of Additively Manufactured Orthopaedic Implants There is now more liberty in the designing aspect due to the growing prominence of AM in this field. Despite the advantages these two processes have, they also have their limitations such as the limited materials that can be processed by them as well as the restriction of using single material per build. Furtherm ore, the dimensional accuracy of the designed cellular structures and the entrapment of powder within them for orthopaedic implants proved to be challenging. However, these limitations can be overcome in the near future by further improvement in the SLM and EBM systems such as reduction of the laser or electron beam spot size for higher accuracy, detailed research in laser or electron beam interaction with powders to minimize powder adhesions. Methods to remove the trapped powders can also be designed such as chemical etching, ultrasonic vibration and post processing treatments. In conclusion, the review carried out in this work showed that SLM and EBM has great potential in the field of orthopaedic applications, however, they have not been widely adopted yet because of some of the limitations listed above. Nonetheless, in the long run, SLM and EBM provides faster and more flexible manufacturing routes compared to conventional manufacturing methods. With these advantages they provide, exciting researches will emerge to take advantage of such technologies in this field.
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AUTHORS’ CONTRIBUTIONS Mr. Swee Leong Sing is the main author of the paper. His contribution includes the research for literature for the bulk of this review paper, inclusive of describing selective laser melting and electron beam melting, design aspects, material properties and biological response, as well as the overall outline and flow of the paper. Dr. Jia An contributes to the biological aspects in the paper and link them to the designs and material properties. Dr. Wai Yee Yeong and Dr. Florencia Edith Wiria contribute to various sections of the paper for material consistency and language. They also ensures that the paper is concise and easy to digest for readers. The authors hereby confirm that this manuscript is their original work and has not been published nor has it been submitted simultaneously elsewhere. Moreover, they confirm that all authors have checked the manuscript and have agreed to the submission.
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