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Additive manufacturing (AM)
Additive manufacturing (AM)
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Mohammadreza Namatollahi*, Ahmadreza Jahadakbar*, Mohammad Javad Mahtabi*,†, Mohammad Elahinia* *Dynamic and Smart Systems Laboratory, Mechanical Industrial and Manufacturing Engineering Department, The University of Toledo, Toledo, OH, United States, †Department of Mechanical Engineering, University of Tennessee at Chattanooga, Chattanooga, TN, United States
12.1
Introduction
Additive manufacturing (AM), also known as three-dimensional (3D) printing and rapid prototyping, is the technology of making solid objects through layer-by-layer addition of material based on a computer-aided design (CAD). Before the invention of AM, most of the metallic parts, including medical implants, were fabricated using traditional processes such as forging, investment casting, hot rolling, and machining. The AM is a powerful and flexible fabrication technique that has gained significant popularity in various applications, from aerospace to biomedical industries. In recent years, AM techniques have been developed to a level that they can realize fabrication of parts with desired functionality. In medical devices, AM can be of great interest due to a high demand for custom-made (i.e., patient-specific) medical devices and biomedical implants with complex geometries. Moreover, the precision of the final product is very important in medical parts. The process of developing a medical device using AM techniques involves various steps, as illustrated in Fig. 12.1. After determining the requirement from AM, imaging and scanning may be needed to determine the geometry of the part that is to be fabricated, unless a completely new design is needed. This imaging and scanning can be performed using advanced techniques, such as computed tomography (CT scan). Based on the imaging followed by necessary customizations, a 3D model needs to be constructed as the input to the computer program that controls the 3D printer. Sometimes, more advanced computer simulation, such as finite element analysis (FEA), is needed of the designed model to assess the biomechanical properties or to design a surgery planning. Then, the 3D geometry will be sliced using a computer program to be used by a 3D printing software and manufacturing/processing parameters are assigned for the fabrication. The processing parameters depend on the method of AM and include laser power, laser scanning speed, layer thickness, scanning pattern, hatch spacing (i.e., distance between two adjacent passes of laser). After fabrication, based on the “indications for use” of the fabricated part, different regulatory approvals and tests are required to evaluate the biological considerations including biocompatibility, cleaning, and sterility.
Metals for Biomedical Devices. https://doi.org/10.1016/B978-0-08-102666-3.00012-2 © 2019 Elsevier Ltd. All rights reserved.
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Fig. 12.1 Process of design and fabrication of AM medical devices ( Javaid and Haleem, 2017).
In this chapter, various methods of AM, which have been used to fabricate metallic medical devices, are introduced followed by the advantages and limitations of AM. Next, different applications that can benefit from AM are introduced and some examples are provided. Statistics about the AM medical industry and the associated market are also provided. Biological considerations for a successful use of AM for medical devices are discussed next. Finally, a case study is presented on an additively manufactured bone fixation plate made of NiTi alloys.
12.2
AM techniques for medical devices
Several AM technologies have been developed in the past two decades for free-form fabrication of different materials: stereolithography for photocurable resins, laminated object manufacturing for any material in sheet form, material jetting for polymers and waxes, binder jetting, material extrusion and powder bed fusion for various materials, to name a few. Among the AM techniques, selective laser melting (SLM), direct energy deposition (DED), and electron beam melting (EBM) are the most common
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techniques for the fabrication of metallic medical parts. In all of these methods, first a 3D CAD of the product is needed. As mentioned before, the CAD model can be designed or prepared based on the data from 3D scanners or CT scans. Then the CAD file will be sliced into subsequent two-dimensional (2D) layers using appropriate software. Finally, the desired part is printed layer by layer using one of the aforementioned techniques. A brief description of each technology will be discussed in the following sections.
12.2.1 Selective laser melting The SLM is a laser powder bed fusion (LPBF) technique that uses a laser beam to melt a powder bed selectively. For SLM AM of metals, the material should be in the form of fine powder. The powder can be generated from an ingot using various techniques, such as gas atomization or plasma atomization techniques. Fig. 12.2 shows a schematic illustration of an SLM AM process. To start the fabrication, this fine powder is deposited on a substrate using a roller or scraper. The thickness of the powder layer is usually 20–100 μm. Then a laser beam melts the desired 2D cross section of the first layer (selective melting), based on the sliced CAD file, resulting in melting and fusing of those areas. Afterwards, the substrate moves down by the thickness of each layer, which is set by the user, and a new layer of powder is deposited. This layer-by-layer process will continue until the part is completely fabricated. Laser power, scanning speed, hatch spacing, and powder bed layer thickness are some of the most important parameters that are involved in SLM process. These parameters need to be optimized, for each material, to achieve a successful fabrication. The process parameters directly affect the microstructure, density, mechanical behavior, strength, and surface quality of the final product. Due to high reactivity of materials
X-Y scanner system
Laser Laser beam Roller
Part
Powder bed
Base plate
Over flow container
Powder delivery system Z Fabrication piston
Fig. 12.2 Schematic illustration of an SLM AM process (Shamsaei et al., 2015).
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in high temperatures, usually the SLM process is carried out in an inert atmosphere, such as argon, to lower the impurity pick-up level and produce medical grade products. The SLM enables the fabrication of metallic parts with excellent mechanical properties, because of the characteristics of grain refinement, extended solid solubility, chemical homogeneity, and reduction in quantity and size of phase segregation.
12.2.2 Direct energy deposition The direct laser deposition (DLD) is another type of metal AM method in which instead of having a powder bed, the powder is selectively deposited and simultaneously melted by a laser beam, resulting in a molten pool on the substrate. The laser/deposition head and/or the substrate can travel in space directions, demonstrating the ability to fabricate 3D objects. Fig. 12.3 schematically demonstrates the DED AM technique. As shown in this figure, the powder spray nozzles are located next to the laser and can move with laser, if needed. The number of the nozzles may vary from machine to machine and a DED machine can be equipped with single or multiple nozzles. The most important process parameters in this method that need to be optimized are laser-substrate relative velocity (traverse speed), laser power and beam diameter, hatch spacing, powder feeding rate, and scanning strategy. Thanks to the control flexibilities of the laser beam and powder deposition, DED can be used to repair existing parts or add new structures to the currently fabricated parts.
Turning mirror Laser
Pyrometer
Deposition head Inert gas Lens Powder delivery nozzle
Infrared camera
Nozzle insert Powder stream Focal plane Focused laser beam
Melt pool Deposited metal Fixed substrate
Fig. 12.3 Schematic illustration of direct energy deposition (DED) additive manufacturing technique (Shamsaei et al., 2015).
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Fig. 12.4 Electron beam melting schematics (Galati and Iuliano, 2018).
12.2.3 Electron beam melting Fig. 12.4 shows the features of EBM technique. The heat source in EBM is an electron beam that melts the powder. Similar to SLM the EBM parts are fabricated layer by layer but in a high vacuum chamber, which makes this method suitable for highly reactive materials, such as titanium (Ti) and alloys. Another advantage of EBM over other AM techniques is that the process occurs at temperatures as high as 1000°C resulting in low residual stresses during fabrication and eliminating the need of post heat treatment. A comparison between capabilities and resulting properties of biomedical components fabricated via different AM techniques, explained above, is presented in Table 12.1 Also, Table 12.2 provides a summary of the laser type, fabrication resolution, and the type of materials that can be used in each of the discussed AM techniques.
12.3
Advantages, limitations, and challenges of AM
Similar to any other fabrication technique, there are advantages and challenges associated with using AM in medical devices. In this sections some of the main advantages and challenges associated with these technologies are discussed, which may help in making a decision in using AM for certain medical applications.
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Table 12.1 Comparison between the various AM techniques for fabricating biomedical components (Trevisan et al., 2017). Method
Component
Mechanical properties
Ref.
SLM
Porous Ti-6Al-4V scaffolds Porous Ti-6Al-4V structures Porous Ti structures
E ¼ 3.7–6.7 Pa UCS ¼ 145–164 MPa E ¼ 7–60 GPa YS ¼ 471–809 MPa E ¼ 2.6–44 GPa YS ¼ 24–463 MPa E ¼ 3.9–12.9 GPa YS ¼ 49.6–107.5 MPa UCS ¼ 127.1–148.4 MPa E ¼ 14.5–38.5 GPa YS ¼ 138–194 MPa UCS ¼ 163–286 MPa
Wieding et al. (2012) Bandyopadhyay et al. (2010) Xue et al. (2007)
DED DED EBM
Compact and porous Ti-6Al-4V structures
EBM
Porous Ti-6Al-4V implant
Ponader et al. (2009) Li et al. (2012)
E, modulus of elasticity; UCS, ultimate compressive stress; YS, yield stress.
Table 12.2 Common materials, resolution, and laser type for each of the mentioned powderbased metal AM methods (Shayesteh Moghaddam et al., 2017). Method
Common materials
Resolution
Laser type
SLM
Stainless steels, aluminum (Al), copper (Cu), iron (Fe), cobalt (Co)chrome (CR), Ti, Ni-based alloy, and a mixture of different types of particles (Fe, Ni, Cu, and Fe3P) (Abe et al., 2001; Childs et al., 2005; Kruth et al., 2004) Steel alloys, stainless steel, tool steel, Al, bronze, Co-Cr, (Fe, Ni)TiC composites and Ti (Xiao and Zhang, 2007) Ti, Ti-6Al-4V, Ti Grade 2, Co-Cr
20 μm (Ravari et al., 2015)
Nd:YAG (Delgado et al., 2012)
20 μm (H€anninen, 2001)
CO2 (Zhu et al., 2003), Fiber (Schleifenbaum et al., 2010) N/A
DED
EBM
N/A
12.3.1 Advantages of AM The advantages of AM in medical device industry originate from both the needs of the medical industry—such as complex geometries, patient-to-patient variation, high required precision, etc.—as well as the capabilities of AM, such as flexibility in geometry and fabrication of complex geometries, such as porous lattice structures, that conventional methods are unable to make. Using AM, completely customized parts can be fabricated that can be very advantageous for medical devices, where the device may vary from patient to patient. Moreover, the use of different materials to fabricate medical devices and parts is easily possible in AM, as long as the material can be produced
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to a form that can be used by AM machines. For an optimized AM process, very little postfabrication treatment is needed. For AM, a 3D model of the final product is generated in a computer before fabrication. This provides the opportunity of virtually assessing the appropriateness of the product and its match with the final implantation with the anatomy of the patient. In such cases, modifications can be made to the product quickly, before making the device. Finally, because of all these characteristics, AM of medical devices has a noticeably shorter supply-chain lead time and, in general, lower production cost.
12.3.2 Challenges of AM Although AM technologies are advantageous in making medical devices, there are challenges toward achieving a successful fabrication with desired functional and structural properties. Indeed, first time fabrication of parts with a material could be challenging until a set of appropriate processing parameters (e.g., laser power, scanning speed, etc., in a laser-based manufacturing) could be determined. This procedure adds to the lead time and cost of fabrication in an AM process. On the other hand, since AM is in its early stages of development, there are a lot of uncertainties on the effects of various factors, such as shape and size of the part, on the properties of the fabricated parts. Understanding and tailoring the influence of various fabrication parameters on the mechanical properties of the part requires extensive research. This also requires regulations for fabricating and characterizing AM parts to ensure consistency of fabrications using different machines and at different fabrication sites. Dimensional accuracy of the AM parts is another challenge with fabricating metallic medical devices especially for miniature components. In addition, providing an acceptable surface quality of the fabricated part, to eliminate postfabrication modifications, requires a lot of work and may not be possible for most materials and geometries at the current state of metal AM. Finally, durability of the fabricated AM parts is still the main challenge toward using these parts for medical applications, as majority of the real-world applications involve cyclic loading applied to these components. At the current state of AM, imperfections such as voids and impurities are inevitable in a fabricated AM component. Moreover, surface finish of the as-fabricated AM parts shows rough features. All these fabrication imperfections contribute to the fatigue failure of the AM devices and reduce their durability (Bagheri et al., 2018; Mahtabi et al., 2015a; Yadollahi et al., 2018).
12.4
Applications of AM in medical devices
The applications of AM in various industries has grown rapidly during the past decade. Especially in medical industry, these applications range from educational aspects to fabrication of final medical devices and implants ( Javaid and Haleem, 2017; Tuomi et al., 2014; Zadpoor and Malda, 2017). While applications of AM in biomedical applications can be categorized based on different perspectives, some of the major application of AM in medical devices are discussed in this section.
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Fig. 12.5 A prototype forceps fabricated using additive manufacturing (Kontio et al., 2012).
12.4.1 Tools, guides, and instruments for medical devices One of the applications of AM technology in medical devices is to fabricate tools and instruments to improve the effectiveness of a medical procedure or surgery. The main advantage of AM for this type of application is that the complete process from design to fabrication can be done at once and the production process is much faster and more consistent. AM drill guides (Bibb et al., 2009; Labadie et al., 2008), oral appliances (Salmi et al., 2013), and surgical instruments/guides (Kontio et al., 2012) are examples of such applications. Fig. 12.5 shows an example of the additively manufactured tools and instruments. Using AM, instruments can be made for a specific operation—which are not available in the market—to meet the especial needs of the operation. As the devices and instruments for this type of application are in direct contact with body fluids and tissues for a limited time, one important consideration is that the material used for this type of application must be sterilizable, but does not need to be implantable (Tuomi et al., 2014).
12.4.2 Prosthesis and orthotics The AM has been also used in orthotic and prosthetic applications. Similar to other applications, these applications benefit from the capability of AM in creating customized parts. Moreover, for prosthetic devices, AM enables fabricating functionally tuned devices by matching the dimensions of the device. For such applications, multimaterial technologies that are able to print colorful materials, which are not necessarily metals, are very advantageous. An additively manufactured prosthetic hand is shown in Fig. 12.6.
12.4.3 Customized implants Freedom in form is the main advantage of AM for fabricating customized, patientspecific implants. Using this capability, implants can be fabricated to match the anatomy of a patient. Examples of implants that are fabricated using various AM techniques are presented in Fig. 12.7 Implant applications of AM are so advanced that
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Fig. 12.6 Examples of AM orthotic and prosthetic parts: (A) Prosthetic hand (Burn et al., 2016), (B) foot orthosis (left) and ankle-foot orthosis (right) (Chen et al., 2016).
there are already companies that offer patient-specific AM implants following the procedure explained in Section 12.2. In addition, complex geometries such as lattice structures with intentional porous structures can be fabricated using different AM techniques. Examples of such lattice structures are shown in Fig. 12.7B. The lattice structures can be designed to meet different requirements, such as matching the mechanical properties of the implant to the biological tissue, and preventing some postimplantation issues, such as stress shielding. The porous structure of the implants also provides the potential for growth of tissue during healing, as well as osseointegration. The AM was used to fabricate artificial teeth as well (Liang et al., 2018) (Fig. 12.7C).
12.4.4 Education and training Another important use of AM is its use for educational and training purposes, as well as for communicating with patients. In addition, 3D reconstruction of the devices and physical parts (Zadpoor and Malda, 2017; Spottiswoode et al., 2013; Zein Nizar et al., 2013) using an AM can be very helpful for planning complex surgeries. The
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(A)
(C)
(B)
(D)
Fig. 12.7 Examples of medical implants fabricated by AM techniques: (A) Ti alloy dental implant (Obaton et al., 2017), (B) NiTi bone fixation plate, (C) lattice structure of NiTi fabricated by SLM technique, and (D) as-fabricated titanium sternum/rib implant made by EBM technique (Thompson, 2018).
capabilities of AM can also help evaluate the implantation procedure and the match between patient’s anatomy and the medical device. The AM parts can be fabricated to represent human anatomy and for educating new surgeons and medical students by providing a more realistic environment. Multimaterial printing can be very helpful in this type of application by providing colorful fabricated objects.
12.4.5 Drug delivery In 2015, food and drug administration (FDA) approved drug products fabricated by AM (Goyanes et al., 2016), indicating the great potential of AM for drug delivery devices in the solid form (Zadpoor and Malda, 2017). The main AM techniques that can be used for this type of application are binder jetting techniques, material extrusion, and material jetting. Using an AM, the drug release profile can be controlled by
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Fig. 12.8 Revenue split of AM equipment customers (Europe, 2015).
adjusting the shape of the device. Moreover, spatial distribution of the active agent in the device can also be designed in AM devices to have more control over the release profile. Another advantage of AM for drug delivery devices is the capability of fabricating multidrug materials into a single delivery device.
12.5
Market and statistics
The AM market is one of the world’s fastest growing markets with a predicted market value of $10.8 billion by 2021. Fig. 12.8 shows the revenue of AM medical devices (Europe, 2015). It can be seen that the most rapid growth is in automotive, electronics, and medical industries. Furthermore, considering very broad and generic applications as well as increasing number of accidents causing injuries and diseases there are demands in the market for faster AM of medical devices. The advantage of designing patient-specific devices is,
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another reason toward additively manufactured medical devices in the market, to such an extent that it is estimated that in the near future AM will be used for 80% of the implants. Reports show that a total market of approximately $1.5 billion is anticipated for AM of medical devices by 2026 (Liang et al., 2018).
12.6
Biological considerations and biocompatibility of AM parts
12.6.1 Biological consideration Technological advancement and availability of AM methods have led to increased investment and use by medical device industry. In December 2017, the United States FDA issued a comprehensive guidance for industry and FDA staff on “Technical Considerations for Additive Manufactured Medical Devices” (Food and Drug Administration, 2017). Beside other technical and mechanical considerations, this FDA document introduces and discusses all the parameters and conditions for the AM process that can affect the biocompatibility of the fabricated parts. While different AM techniques have been used in medical device industry, according to FDA—in addition to biocompatibility—other critical considerations, including cleaning and sterility, must be evaluated for AM medical devices. In general, achieving adequate cleaning, sterility, and biocompatibility is a major concern for different methods of AM and can be a limiting factor for the fabrication of specific geometries and materials. For instance, porous or internally complex devices must be carefully designed and processed due to the challenging procedure of cleaning and sterilization of these parts. In this section first we discuss the general aspects of biological considerations of AM devices, including cleaning, sterility, and biocompatibility, and then as an example, the biological consideration of additively manufactured Ti-6Al-4V will be discussed in more detail. For all methods of AM fabrication, each of the process parameters including the raw material, building atmosphere, energy input properties, as well as any postfabrication treatments may significantly affect the biological properties of the final part. The first item to consider is the material composition of the final AM part after postfabrication treatments. Process parameters such as the energy input and the existing elements in the build chamber may cause unexpected chemical reactions, which may adversely affect the biocompatibility of the fabricated parts. If this is the case, the content and build atmosphere properties (i.e., temperature, humidity, etc.) of the AM process must be controlled and monitored during the fabrication. In some cases, an inert gas such as argon is fed into the built chamber to decrease the oxygen level to specific levels to prevent the oxidation of the main material. Postfabrication processing, especially heat treatment, may have the same effect and cause the formation of a nonbiocompatible element. Although AM facilitates the creation of devices with complex geometries, such as components with engineered porosity, internal voids or cavities that may not exist in traditional manufacturing methods, are currently inevitable in AM products.
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Therefore, one should also consider the increased difficulty in removing manufacturing material residues (cleaning) and in the sterilization due to the likelihood of increased surface area, generation of extensive tortuous pathways, and creation of internal voids with limited or no access. In cases where the residual materials could have an adverse effect on the biological aspect of the fabricated part, the manufacturer must establish and maintain procedures for the use and removal of such residual manufacturing material to ensure that it is removed or limited to an amount that does affect the part’s features. For complex AM fabricated geometries, it should be mentioned that since many end-user facilities may not have access to equipment or materials required to implement specific cleaning procedures, only devices that are sufficiently clean of residual manufacturing materials should be provided to the end user. To evaluate the biocompatibility of the final finished devices, the same procedure as in conventional devices must be considered. For any medical device, FDA suggests to evaluate the biocompatibility as described in the guidance “Use of International Standard ISO-10993, ‘Biological Evaluation of Medical Devices Part 1: Evaluation and Testing within a Risk Management Process’” (Anand, 2000). Based on the FDA document, the first step for the assessment of the biocompatibility is the risk management for biocompatibility evaluations. One should identify and asset the potential risks based on the application of the medical device. For instance, a medical device that is in temporary contact with skin needs different types of evaluations compared to a permanent cardiovascular implant. After the risk assessment, a series of required tests should be designed and reported. These tests could be one or a couple of biocompatibility tests including, cytotoxicity, sensitization, hemocompatibility, pyrogenicity, implantation, genotoxicity, carcinogenicity, reproductive and developmental toxicity, and degradation assessments.
12.6.2 Biocompatibility of AM Ti and its alloys Ti and its alloys are among the most well-known biocompatible metals that can be fabricated via AM. In general, in addition to the biocompatibility, bioadhesion, biofunctionality, and corrosion resistance are the other important properties for biomedical applications. The biocompatibility of Ti is mainly due to its low electrical conductivity, which contributes to the electrochemical oxidation of Ti leading to the formation of a thin passive oxide layer on its surface (Quinn and Armstrong, 1978). This oxide layer creates a high resistance to corrosion. In addition, this protective oxide layer remains at pH values of the human body (Schiff et al., 2002). In this section, as an example, some of the studies on biocompatibility of AM Ti and its alloys are discussed in more details. Commercially pure Ti (CP-Ti) and Ti-6Al-4V are the most well-known compositions for the fabrication of Ti implants. The American Society for Testing and Materials (ASTM F67-06, 2006) specifies CP-Ti and Ti-6Al-4V as grade 1–5. Grades 1–4 are unalloyed CP-Ti and grade 5 is the Ti-6Al-4V alloy (ASTM F136-08, 2008). Table 12.3 summarizes the mechanical properties and different grades of Ti based on ASTM standards F67 and F136 (ASTM F67-06, 2006, 2008). The first step to
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Table 12.3 Selected mechanical requirements properties for an implantable Ti Bar (Elias et al., 2008).
Material
Specification
CP-Ti
ASTM F67 Grade 1 ASTM F67 Grade 2 ASTM F67 Grade 3 ASTM F67 Grade 4 ASTM F67 Grade 5
CP-Ti CP-Ti CP-Ti Ti-6Al4V
Tensile strength (MPa)
0.2% proof stress (MPa)
Elongation (%)
Elastic modulus (GPa)
240
170
24
103–107
345
275
20
109–107
450
380
18
103–107
550
483
15
103–107
860
795
10
114–120
fabricate a biocompatible Ti is to reach to the standard mechanical properties based on the standards. In addition to the mechanical properties, the chemical composition of the fabricated part must fulfill some standard requirements. ASTM and ISO have recently published new standards (ISO 22068-14, 2014; ASTM F3001-14, 2014; ASTM F2924-14, 2014; ASTM 2885-11, 2011) for additively layered manufactured (ALM) Ti parts, which are summarized in Table 12.4. These standards also include the minimum mechanical properties of the fabricated parts via ALM and is summarized in Table 12.5. Following the mechanical properties and chemical composition, in vitro and in vivo studies have been performed on the AM fabricated Ti parts that verify the biocompatibility and reliability of this material. These studies used cultured fibroblasts Table 12.4 Chemical composition requirements for ALM Ti according to ASTM standards. Chemical composition (wt%) Specification
Al
V
Fe
C
N
O
H
Y
ASTM F2885 Grade 5 ASTM F2924 ALM Ti-6Al4V ASTM F3001 ALM Ti-6Al4V
5.5–6.75
3.5–4.5
<0.30
<0.08
<0.05
<0.20
<0.015
<0.005
5.5–6.5
3.5–4.5
<0.3
<0.08
<0.05
<0.20
<0.015
–
5.5–6.5
3.5–4.5
<0.25
<0.08
<0.05
<0.13
<0.012
<0.005
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Table 12.5 Mechanical properties of ALM Ti-6Al-4V.
Material
Specification
Ti-6Al4V ALM Ti-6Al4V ALM Ti-6Al4V
ASTM F2885 Grade 5 ASTM F2924 ALM Ti-6Al4V ASTM F3001 ALM Ti-6Al4V (ELI) Tensile
Tensile strength
0.2% proof stress
Elongation (%)
Reduction of area
780
680
10
15
895
825
10
15
860
795
8
25
and osteoblasts in the observation of cell responses to surfaces and also human and animal subjects. Table 12.6 summarizes the processing methods, test types, and general findings of these studies.
12.7
Case study: Stiffness-matched bone fixation plates for mandibular reconstruction surgery
Mandibular segmental defects usually happen due to tumors, cancers, radiation therapy, and trauma. The standard of care for the treatment of a segmental defect starts with resectioning the defected region and using a bone graft (i.e., a healthy piece of bone harvested from patient’s own body, such as fibula) to fill the resected region. As a part of this procedure, mechanical fixations including metallic bone plates and screws are used to immobilize the pieces of bone, and to facilitate and expedite the healing process (Fig. 12.9A) (Shayesteh Moghaddam et al., 2016). Conventionally, these bone plates are made of Ti-6Al-4V alloys. Although biocompatible Ti-6Al4V bone plates provide enough immobilization immediately after the surgery, they may cause major complexities when the bones heal. Due to the high stiffness of Ti-6Al-4V in comparison to cortical bones, the load and stress distribution on the surrounding bones is affected in a way that a major portion of the loads are carried by the fixation plates instead of the bones. This reduced level of stress, which is also called stress shielding, leads to bone resorption and eventually failure of the surgery based on the bone remodeling phenomena (Mertens et al., 2014). The existence of bone resorption due to stress shielding can be clearly seen in the hip implant, thanks to X-ray images, in Fig. 12.9B. As it can be seen in this figure, the surrounding bones, especially on the top right portion of the hip implant, are darker in comparison with the healthy case that shows the bone resorption due to the stress-shielding effect. One solution to prevent this failure is to remove the bone plates after the bonehealing period in a secondary surgery. Although removing the bone plates could
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Table 12.6 Summary of AM processing methods and findings. Method
Alloy
Biocompatibility test
Cell line/ implantation
EBM
Ti-6Al-4V
In vitro
Human fetal osteoblasts (hFOB 1.19)
EBM
Ti-6Al-4V
In vivo
Frontal skull of domestic pig
EBM
Ti-6Al-4V
In vivo
Rabbit femur and tibia
EBM
Ti-6Al-4V
In vitro
Human adiposederived adult stem cells (hASC)
EBM
Ti-6Al-4V
In vitro and in vivo
EBM
Ti-6Al-4V
In vivo
Osteoblasts extracted from Calvaria of rabbits, Calvaria of rabbits Sheep
DMLS
Ti-6Al-4V
In vitro
Human osteoblasts cells (HOB)
DMLS
Unspecified
In vitro and in vivo
DMLS
Ti-6Al-4V
In vivo
BMP-7 transduced human gingival Fibroblasts Human anterior mandible, minipig mandibular
General findings Reduced cell proliferation in highly rough surfaces (ASTM F136-13, 2013) More bone contact in more porous samples (Ponader et al., 2009) As-EBM implant response comparable to machined (Thomsen et al., 2008) Increased proliferation on porous compared to polished and unpolished EBM discs (Haslauer et al., 2010) Proliferation in porous EBM Ti-6Al4V implants matched coated implants (Li et al., 2012) High bone-implant contact in porous implant (Palmquist et al., 2011) Cultured cells attached and proliferated on SLM substrates (Hollander et al., 2003) In vitro and in vivo test data showing substantial bone ingrowth (Hollister et al., 2005) Peri-implant bone in close contact with the surface of the implant (Mangano et al., 2010)
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Table 12.6 Continued Method
Alloy
Biocompatibility test
Cell line/ implantation
DMLS
Ti-6Al-4V
In vitro
Human osteoblasts human
LENS
CP-Ti
In vitro
LENS
Ti-6Al-4V
In vivo
Human osteoblast cells (OPC1) Male SpragueDawley rats
Modified FDM
CP-Ti
In vitro
L929 mouse fibroblast
General findings Osteoblasts wellspread and with multiple contact points (Warnke et al., 2008) Cells well spread on porous Ti (Xue et al., 2007) Increase in calcium (bone) within implant pores (Bandyopadhyay et al., 2010) Excellent bone cell attachment and proliferation (Wiria et al., 2010)
Fig. 12.9 (A) Schematic illustration of a mandibular reconstruction ( Jahadakbar et al., 2016) and (B) the effect of stress shielding and bone resorption (Engh et al., 1987).
prevent the risk of failure, extra cost and risk of the second surgery, and the increased recovery time are the main side effects of this solution (Boyd et al., 2012). Recently, patient-specific, stiffness-matched bone plates have been introduced to prevent the stress-shielding phenomenon and to eliminate the need for a second surgery ( Jahadakbar et al., 2016). These bone plates are made of Ni-Ti alloy (NiTi or Nitinol), designed based on medical images of the patient (CT scans or X-rays), and are fabricated via AM. The stiffness-matching option minimizes the stress shielding on the
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Fig. 12.10 The process of design and fabrication of patient-specific, stiffness-matched mandibular bone fixation plates.
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surrounding bone and prevent bone resorption. Therefore, there is no need to remove the bone plates with a second surgery. It is worth noting that NiTi, which is the base material for this next generation of bone plates, is a biocompatible shape memory alloy with unique features such as superelasticity and high fatigue life at large levels of strains, which make it a great candidate for biomedical implants (Mahtabi et al., 2015b,c; Mahtabi and Shamsaei, 2017). Due to the high thermomechanical sensitivity of this alloy, the conventional methods of fabrication such as machining and forging are very challenging for this material (Elahinia et al., 2012). Due to these manufacturing limitations, in the past NiTi was only available in the form of wire, rod, tube, and sheet. Recently, AM has been used for successful fabrication of this alloy with almost no limitations directly from CAD files (Elahinia et al., 2016). The mechanical properties and biological properties of the AM fabricated NiTi have been compared with conventionally fabricated NiTi and confirmed the successful fabrication of this alloy via SLM technique (Habijan et al., 2013; Haberland et al., 2012; Ibrahim et al., 2018). Fig. 12.10 shows a general schematic of the process of AM of mandibular fixation plate. The design of these bone plates starts with CT-scan data of the patient’s mandible. Based on the CT-scan data, first a 3D CAD model of the patient’s mandible is created. Then in a virtual process the defected region is resected and replaced with a double barrel fibula graft. Following this step, the general geometry of the required bone plates is designed. Then the reconstructed mandible along with the bone plates and screws are transferred to a finite element (FE) software and modeled. In a series of simulations, using the CT-scan data, the required stiffness of the bone plates is calculated. Next, specific levels and types of porosity are assigned to the bone plates to modulate the final stiffness of the bone plates to the required level. Finally the CAD file of the patient-specific, stiffness-matched bone plates are designed based the patient’s anatomy and FE simulations is used for AM ( Jahadakbar et al., 2016).
References Abe, F., Osakada, K., Shiomi, M., Uematsu, K., Matsumoto, M., 2001. The manufacturing of hard tools from metallic powders by selective laser melting. J. Mater. Process. Technol. 111 (1–3), 210–213. Anand, V.P., 2000. Biocompatibility safety assessment of medical devices: FDA/ISO and Japanese standards. Med. Device Diagn. Ind. 22 (1), 206–219. ASTM 2885-11, 2011. Standard Specification for Metal Injection Molded Titanium-6Aluminum-4Vanadium Components for Surgical Implant Applications. ASTM International, West Conshohocken, PA. ASTM F136-08, 2008. Standard Specification for Wrought Titanium-6 Aluminum-4 Vanadium ELI (Extra Low Interstitial) Alloy for Surgical Implant Applications (UNS R56401). ASTM International, West Conshohocken, PA. ASTM F136-13, 2013. Standard Specification for Wrought Titanium-6 Aluminum-4 Vanadium ELI (Extra Low Interstitial) Alloy for Surgical Implant Applications (UNS R56401). ASTM International, West Conshohocken, PA.
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ASTM F2924-14, 2014. Standard Specification for Additive Manufacturing Titanium-6 Aluminum-4 Vanadium ELI with Powder Bed Fusion. ASTM International, West Conshohocken, PA. ASTM F3001-14, 2014. Standard Specification for Additive Manufacturing Titanium-6 Aluminum-4 Vanadium ELI (Extra Low Interstitial) with Powder Bed Fusion. ASTM International, West Conshohocken, PA. ASTM F67-06, 2006. Standard Specification for Unalloyed Titanium, for Surgical Implant Applications (UNS R50250, UNS R50400, UNS R50550, UNS R50700). ASTM International, West Conshohocken, PA. Bagheri, A., Mahtabi, M.J., Shamsaei, N., 2018. Fatigue behavior and cyclic deformation of additive manufactured NiTi. J. Mater. Process. Technol. 252, 440–453. Bandyopadhyay, A., Espana, F., Balla, V.K., Bose, S., Ohgami, Y., Davies, N.M., 2010. Influence of porosity on mechanical properties and in vivo response of Ti6Al4V implants. Acta Biomater. 6 (4), 1640–1648. Bibb, R., Bocca, A., Eggbeer, D., Evans, P., Sugar, A., 2009. Rapid manufacture of customfitting surgical guides. Rapid Prototyp. J. 15 (5), 346–354. Boyd, T.G., Huber, K.M., Verbist, D.E., Bumpous, J.M., Wilhelmi, B.J., 2012. CASE REPORT removal of exposed titanium reconstruction plate after mandibular reconstruction with a free fibula osteocutaneous flap with large surgical pin cutters: a case report and literature review. Eplasty 12, e42. Burn, M.B., Ta, A., Gogola, G.R., 2016. Three-dimensional printing of prosthetic hands for children. J. Hand Surg. Am. 41 (5), e103–e109. Chen, R.K., Jin, Y.-a., Wensman, J., Shih, A., 2016. Additive manufacturing of custom orthoses and prostheses—a review. Addit. Manuf. 12, 77–89. Childs, T., Hauser, C., Badrossamay, M., 2005. Selective laser sintering (melting) of stainless and tool steel powders: experiments and modelling. Proc. Inst. Mech. Eng. B J. Eng. Manuf. 219 (4), 339–357. Delgado, J., Ciurana, J., Rodrı´guez, C.A., 2012. Influence of process parameters on part quality and mechanical properties for DMLS and SLM with iron-based materials. Int. J. Adv. Manuf. Technol. 60 (5–8), 601–610. Elahinia, M.H., Hashemi, M., Tabesh, M., Bhaduri, S.B., 2012. Manufacturing and processing of NiTi implants: a review. Prog. Mater. Sci. 57, 911–946. Elahinia, M., Shayesteh Moghaddam, N., Taheri Andani, M., Amerinatanzi, A., Bimber, B.A., Hamilton, R.F., 2016. Fabrication of NiTi through additive manufacturing: a review. Prog. Mater. Sci. 83, 630–663. Elias, C.N., Lima, J.H.C., Valiev, R., Meyers, M.A., 2008. Biomedical applications of titanium and its alloys. JOM 60 (3), 46–49. Engh, C., Bobyn, J., Glassman, A., 1987. Porous-coated hip replacement. The factors governing bone ingrowth, stress shielding, and clinical results. J. Bone Joint Surg. Br. 69-B (1), 45–55. Europe, M., 2015. An Overview of the Additive Manufacturing Market and Its Role on the Advances in Medical, Bioprinting and Drug Discovery. Medtec Europe. Food and Drug Administration, 2017. Technical Considerations for Additive Manufactured Medical Devices. Food and Drug Administration. Galati, M., Iuliano, L., 2018. A literature review of powder-based electron beam melting focusing on numerical simulations. Addit. Manuf. 19, 1–20. Goyanes, A., Det-Amornrat, U., Wang, J., Basit, A.W., Gaisford, S., 2016. 3D scanning and 3D printing as innovative technologies for fabricating personalized topical drug delivery systems. J. Control. Release 234, 41–48.
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Haberland, C., Meier, H., Frenzel, J., 2012. On the Properties of Ni-Rich NiTi Shape Memory Parts Produced by Selective Laser Melting. pp. 97–104. Habijan, T., Haberland, C., Meier, H., Frenzel, J., Wittsiepe, J., Wuwer, C., Greulich, C., Schildhauer, T.A., K€oller, M., 2013. The biocompatibility of dense and porous Nickel– Titanium produced by selective laser melting. Mater. Sci. Eng. C 33 (1), 419–426. H€anninen, J., 2001. DMLS moves from rapid tooling to rapid manufacturing. Met. Powder Rep. 56 (9), 24–29. Haslauer, C.M., Springer, J.C., Harrysson, O.L.A., Loboa, E.G., Monteiro-Riviere, N.A., Marcellin-Little, D.J., 2010. In vitro biocompatibility of titanium alloy discs made using direct metal fabrication. Med. Eng. Phys. 32 (6), 645–652. Hollander, D.A., Wirtz, T., von Walter, M., Linker, R., Schultheis, A., Paar, O., 2003. Development of individual three-dimensional bone substitutes using “Selective LaserMelting” Eur. J. Trauma 29 (4), 228–234. Hollister, S.J., Lin, C.Y., Saito, E., Lin, C.Y., Schek, R.D., Taboas, J.M., Williams, J.M., Partee, B., Flanagan, C.L., Diggs, A., Wilke, E.N., Van Lenthe, G.H., M€ uller, R., Wirtz, T., Das, S., Feinberg, S.E., Krebsbach, P.H., 2005. Engineering craniofacial scaffolds. Orthod. Craniofac. Res. 8 (3), 162–173. Ibrahim, H., Jahadakbar, A., Dehghan, A., Moghaddam, S.N., Amerinatanzi, A., Elahinia, M., 2018. In vitro corrosion assessment of additively manufactured porous NiTi structures for bone fixation applications. Metals 8 (3), 164. ISO 22068-14, 2014. Sintered-Metal Injection-Molded Materials—Specifications. International Organisation for Standardisation ISO, Geneva. Jahadakbar, A., Shayesteh Moghaddam, N., Amerinatanzi, A., Dean, D., Karaca, E.H., Elahinia, M., 2016. finite element simulation and additive manufacturing of stiffnessmatched NiTi fixation hardware for mandibular reconstruction surgery. Bioengineering 3 (4), 36. Javaid, M., Haleem, A., 2017. Additive manufacturing applications in medical cases: a literature based review. Alex. J. Med. https://doi.org/10.1016/j.ajme.2017.09.003. Kontio, R., Bj€orkstr, R., Salmi, M., Paloheimo, M., Paloheimo, K.-S., Tuomi, J., M€akitie, A.A., 2012. Designing and additive manufacturing a prototype for a novel instrument for mandible fracture reduction. Surgery 3 (2), 2–4. Kruth, J.-P., Froyen, L., Van Vaerenbergh, J., Mercelis, P., Rombouts, M., Lauwers, B., 2004. Selective laser melting of iron-based powder. J. Mater. Process. Technol. 149 (1–3), 616–622. Labadie, R.F., Noble, J.H., Dawant, B.M., Balachandran, R., Majdani, O., Fitzpatrick, J.M., 2008. Clinical validation of percutaneous cochlear implant surgery: initial report. Laryngoscope 118 (6), 1031–1039. Li, X., Feng, Y.-F., Wang, C.-T., Li, G.-C., Lei, W., Zhang, Z.-Y., Wang, L., 2012. Evaluation of biological properties of electron beam melted Ti6Al4V implant with biomimetic coating in vitro and in vivo. PLoS One 7 (12), e52049. Liang, X., Liao, W., Cai, H., Jiang, S., Chen, S., 2018. 3D-printed artificial teeth: accuracy and application in root canal therapy. J. Biomed. Nanotechnol. 14 (8), 1477–1485. Mahtabi, M.J., Shamsaei, N., 2017. Fatigue modeling for superelastic NiTi considering cyclic deformation and load ratio effects. Shape Mem. Superelasticity 3, 250–263. Mahtabi, M.J., Shamsaei, N., Mitchell, M.R., 2015a. Fatigue of Nitinol: the state-of-the-art and ongoing challenges. J. Mech. Behav. Biomed. Mater. 50, 228–254. Mahtabi, M.J., Shamsaei, N., Elahinia, M.H., 2015b. Fatigue of shape memory alloys. In: Shape Memory Alloy Actuators. John Wiley & Sons, Ltd., pp. 155–190.
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Mahtabi, M.J., Shamsaei, N., Rutherford, B., 2015c. Mean strain effects on the fatigue behavior of superelastic Nitinol alloys: an experimental investigation. Procedia Eng. 133, 646–654. Mangano, C., Piattelli, A., d’Avila, S., Iezzi, G., Mangano, F., Onuma, T., Shibli, J.A., 2010. Early human bone response to laser metal sintering surface topography: a histologic report. J. Oral Implantol. 36 (2), 91–96. Mertens, C., Decker, C., Engel, M., Sander, A., Hoffmann, J., Freier, K., 2014. Early bone resorption of free microvascular reanastomized bone grafts for mandibular reconstruction—a comparison of iliac crest and fibula grafts. J. Craniomaxillofac. Surg. 42 (5), e217–e223. Obaton, A.F., Fain, J., Djemaı¨, M., Meinel, D., Leonard, F., Mahe, E., Lecuelle, B., Fouchet, J.J., Bruno, G., 2017. In vivo XCT bone characterization of lattice structured implants fabricated by additive manufacturing. Heliyon 3 (8), e00374. Palmquist, A., Snis, A., Emanuelsson, L., Browne, M., Thomsen, P., 2011. Long-term biocompatibility and osseointegration of electron beam melted, free-form–fabricated solid and porous titanium alloy: experimental studies in sheep. J. Biomater. Appl. 27 (8), 1003–1016. Ponader, S., von Wilmowsky, C., Widenmayer, M., Lutz, R., Heinl, P., K€ orner, C., Singer Robert, F., Nkenke, E., Neukam Friedrich, W., Schlegel Karl, A., 2009. In vivo performance of selective electron beam-melted Ti-6Al-4V structures. J. Biomed. Mater. Res. A 92A (1), 56–62. Quinn, R.K., Armstrong, N.R., 1978. Electrochemical and surface analytical characterization of titanium and titanium hydride thin film electrode oxidation. J. Electrochem. Soc. 125 (11), 1790–1796. Ravari, M.K., Kadkhodaei, M., Ghaei, A., 2015. A microplane constitutive model for shape memory alloys considering tension–compression asymmetry. Smart Mater. Struct.. 24 (7). 075016. Salmi, M., Paloheimo, K.-S., Tuomi, J., Ingman, T., M€akitie, A., 2013. A digital process for additive manufacturing of occlusal splints: a clinical pilot study. J. R. Soc. Interface 10 (84), 20130203. Schiff, N., Grosgogeat, B., Lissac, M., Dalard, F., 2002. Influence of fluoride content and pH on the corrosion resistance of titanium and its alloys. Biomaterials 23 (9), 1995–2002. Schleifenbaum, H., Meiners, W., Wissenbach, K., Hinke, C., 2010. Individualized production by means of high power Selective Laser Melting. CIRP J. Manuf. Sci. Technol. 2 (3), 161–169. Shamsaei, N., Yadollahi, A., Bian, L., Thompson, S.M., 2015. An overview of direct laser deposition for additive manufacturing; Part II: Mechanical behavior, process parameter optimization and control. Addit. Manuf. 8, 12–35. Shayesteh Moghaddam, N., Jahadakbar, A., Amerinatanzi, A., Elahinia, M., Miller, M., Dean, D., 2016. Metallic fixation of mandibular segmental defects: graft immobilization and orofacial functional maintenance. Plast. Reconstr. Surg. Glob. Open 4 (9), e858. Shayesteh Moghaddam, N., Jahadakbar, A., Amerinatanzi, A., Elahinia, M., 2017. Recent advances in laser-based additive manufacturing. In: Bian, L., Shamsaei, N., Usher, J. (Eds.), Laser-Based Additive Manufacturing of Metal Parts: Modeling, Optimization, and Control of Mechanical Properties. CRC Press. Spottiswoode, B.S., van den Heever, D.J., Chang, Y., Engelhardt, S., Du Plessis, S., Nicolls, F., Hartzenberg, H.B., Gretschel, A., 2013. Preoperative three-dimensional model creation of magnetic resonance brain images as a tool to assist neurosurgical planning. Stereotact. Funct. Neurosurg. 91 (3), 162–169.
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Thompson, R.G., 2018. Anatomics 3D-printed titanium implants from head to heel. In: Froes, F.H., Qian, M. (Eds.), Titanium in Medical and Dental Applications. Woodhead Publishing, pp. 225–237 (Chapter 3.2). Thomsen, P., Malmstr€om, J., Emanuelsson, L., Rene, M., Snis, A., 2008. Electron beam-melted, free-form-fabricated titanium alloy implants: material surface characterization and early bone response in rabbits. J. Biomed. Mater. Res. B Appl. Biomater. 90B (1), 35–44. Trevisan, F., Calignano, F., Aversa, A., Marchese, G., Lombardi, M., Biamino, S., Ugues, D., Manfredi, D., 2017. Additive manufacturing of titanium alloys in the biomedical field: processes, properties and applications. J. Appl. Biomater. Funct. Mater. 16 (2), 57–67. Tuomi, J., Paloheimo, K.-S., Vehvil€ainen, J., Bj€orkstrand, R., Salmi, M., Huotilainen, E., Kontio, R., Rouse, S., Gibson, I., M€akitie, A.A., 2014. A novel classification and online platform for planning and documentation of medical applications of additive manufacturing. Surg. Innov. 21 (6), 553–559. Warnke, P.H., Douglas, T., Wollny, P., Sherry, E., Steiner, M., Galonska, S., Becker, S.T., Springer, I.N., Wiltfang, J., Sivananthan, S., 2008. Rapid prototyping: porous titanium alloy scaffolds produced by selective laser melting for bone tissue engineering. Tissue Eng. Part C Methods 15 (2), 115–124. Wieding, J., Jonitz, A., Bader, R., 2012. The effect of structural design on mechanical properties and cellular response of additive manufactured titanium scaffolds. Materials 5 (8), 1336–1347. Wiria, F.E., Shyan, J.Y.M., Lim, P.N., Wen, F.G.C., Yeo, J.F., Cao, T., 2010. Printing of Titanium implant prototype. Mater. Des. 31, S101–S105. Xiao, B., Zhang, Y., 2007. Laser sintering of metal powders on top of sintered layers under multiple-line laser scanning. J. Phys. D Appl. Phys. 40 (21), 6725. Xue, W., Krishna, B.V., Bandyopadhyay, A., Bose, S., 2007. Processing and biocompatibility evaluation of laser processed porous titanium. Acta Biomater. 3 (6), 1007–1018. Yadollahi, A., Mahtabi, M.J., Khalili, A., Doude, H.R., Newman, J.C., 2018. Fatigue life prediction of additively manufactured material: effects of surface roughness, defect size, and shape. Fatigue Fract. Eng. Mater. Struct. 41 (7), 1602–1614. Zadpoor, A.A., Malda, J., 2017. Additive manufacturing of biomaterials, tissues, and organs. Ann. Biomed. Eng. 45 (1), 1–11. Zein Nizar, N., Hanouneh Ibrahim, A., Bishop Paul, D., Samaan, M., Eghtesad, B., Quintini, C., Miller, C., Yerian, L., Klatte, R., 2013. Three-dimensional print of a liver for preoperative planning in living donor liver transplantation. Liver Transpl. 19 (12), 1304–1310. Zhu, H., Lu, L., Fuh, J., 2003. Development and characterization of direct laser sintering Cu-based metal powder. J. Mater. Process. Technol. 140 (1–3), 314–317.
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Mohammadreza Namatollahi*, Ahmadreza Jahadakbar*, Mohammad Javad Mahtabi*,†, Mohammad Elahinia* *Dynamic and Smart Systems Laboratory, Mechanical Industrial and Manufacturing Engineering Department, The University of Toledo, Toledo, OH, United States, †Department of Mechanical Engineering, University of Tennessee at Chattanooga, Chattanooga, TN, United States
12.1
Introduction
Additive manufacturing (AM), also known as three-dimensional (3D) printing and rapid prototyping, is the technology of making solid objects through layer-by-layer addition of material based on a computer-aided design (CAD). Before the invention of AM, most of the metallic parts, including medical implants, were fabricated using traditional processes such as forging, investment casting, hot rolling, and machining. The AM is a powerful and flexible fabrication technique that has gained significant popularity in various applications, from aerospace to biomedical industries. In recent years, AM techniques have been developed to a level that they can realize fabrication of parts with desired functionality. In medical devices, AM can be of great interest due to a high demand for custom-made (i.e., patient-specific) medical devices and biomedical implants with complex geometries. Moreover, the precision of the final product is very important in medical parts. The process of developing a medical device using AM techniques involves various steps, as illustrated in Fig. 12.1. After determining the requirement from AM, imaging and scanning may be needed to determine the geometry of the part that is to be fabricated, unless a completely new design is needed. This imaging and scanning can be performed using advanced techniques, such as computed tomography (CT scan). Based on the imaging followed by necessary customizations, a 3D model needs to be constructed as the input to the computer program that controls the 3D printer. Sometimes, more advanced computer simulation, such as finite element analysis (FEA), is needed of the designed model to assess the biomechanical properties or to design a surgery planning. Then, the 3D geometry will be sliced using a computer program to be used by a 3D printing software and manufacturing/processing parameters are assigned for the fabrication. The processing parameters depend on the method of AM and include laser power, laser scanning speed, layer thickness, scanning pattern, hatch spacing (i.e., distance between two adjacent passes of laser). After fabrication, based on the “indications for use” of the fabricated part, different regulatory approvals and tests are required to evaluate the biological considerations including biocompatibility, cleaning, and sterility.
Metals for Biomedical Devices. https://doi.org/10.1016/B978-0-08-102666-3.00012-2 © 2019 Elsevier Ltd. All rights reserved.
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Fig. 12.1 Process of design and fabrication of AM medical devices ( Javaid and Haleem, 2017).
In this chapter, various methods of AM, which have been used to fabricate metallic medical devices, are introduced followed by the advantages and limitations of AM. Next, different applications that can benefit from AM are introduced and some examples are provided. Statistics about the AM medical industry and the associated market are also provided. Biological considerations for a successful use of AM for medical devices are discussed next. Finally, a case study is presented on an additively manufactured bone fixation plate made of NiTi alloys.
12.2
AM techniques for medical devices
Several AM technologies have been developed in the past two decades for free-form fabrication of different materials: stereolithography for photocurable resins, laminated object manufacturing for any material in sheet form, material jetting for polymers and waxes, binder jetting, material extrusion and powder bed fusion for various materials, to name a few. Among the AM techniques, selective laser melting (SLM), direct energy deposition (DED), and electron beam melting (EBM) are the most common
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techniques for the fabrication of metallic medical parts. In all of these methods, first a 3D CAD of the product is needed. As mentioned before, the CAD model can be designed or prepared based on the data from 3D scanners or CT scans. Then the CAD file will be sliced into subsequent two-dimensional (2D) layers using appropriate software. Finally, the desired part is printed layer by layer using one of the aforementioned techniques. A brief description of each technology will be discussed in the following sections.
12.2.1 Selective laser melting The SLM is a laser powder bed fusion (LPBF) technique that uses a laser beam to melt a powder bed selectively. For SLM AM of metals, the material should be in the form of fine powder. The powder can be generated from an ingot using various techniques, such as gas atomization or plasma atomization techniques. Fig. 12.2 shows a schematic illustration of an SLM AM process. To start the fabrication, this fine powder is deposited on a substrate using a roller or scraper. The thickness of the powder layer is usually 20–100 μm. Then a laser beam melts the desired 2D cross section of the first layer (selective melting), based on the sliced CAD file, resulting in melting and fusing of those areas. Afterwards, the substrate moves down by the thickness of each layer, which is set by the user, and a new layer of powder is deposited. This layer-by-layer process will continue until the part is completely fabricated. Laser power, scanning speed, hatch spacing, and powder bed layer thickness are some of the most important parameters that are involved in SLM process. These parameters need to be optimized, for each material, to achieve a successful fabrication. The process parameters directly affect the microstructure, density, mechanical behavior, strength, and surface quality of the final product. Due to high reactivity of materials
X-Y scanner system
Laser Laser beam Roller
Part
Powder bed
Base plate
Over flow container
Powder delivery system Z Fabrication piston
Fig. 12.2 Schematic illustration of an SLM AM process (Shamsaei et al., 2015).
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in high temperatures, usually the SLM process is carried out in an inert atmosphere, such as argon, to lower the impurity pick-up level and produce medical grade products. The SLM enables the fabrication of metallic parts with excellent mechanical properties, because of the characteristics of grain refinement, extended solid solubility, chemical homogeneity, and reduction in quantity and size of phase segregation.
12.2.2 Direct energy deposition The direct laser deposition (DLD) is another type of metal AM method in which instead of having a powder bed, the powder is selectively deposited and simultaneously melted by a laser beam, resulting in a molten pool on the substrate. The laser/deposition head and/or the substrate can travel in space directions, demonstrating the ability to fabricate 3D objects. Fig. 12.3 schematically demonstrates the DED AM technique. As shown in this figure, the powder spray nozzles are located next to the laser and can move with laser, if needed. The number of the nozzles may vary from machine to machine and a DED machine can be equipped with single or multiple nozzles. The most important process parameters in this method that need to be optimized are laser-substrate relative velocity (traverse speed), laser power and beam diameter, hatch spacing, powder feeding rate, and scanning strategy. Thanks to the control flexibilities of the laser beam and powder deposition, DED can be used to repair existing parts or add new structures to the currently fabricated parts.
Turning mirror Laser
Pyrometer
Deposition head Inert gas Lens Powder delivery nozzle
Infrared camera
Nozzle insert Powder stream Focal plane Focused laser beam
Melt pool Deposited metal Fixed substrate
Fig. 12.3 Schematic illustration of direct energy deposition (DED) additive manufacturing technique (Shamsaei et al., 2015).
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Fig. 12.4 Electron beam melting schematics (Galati and Iuliano, 2018).
12.2.3 Electron beam melting Fig. 12.4 shows the features of EBM technique. The heat source in EBM is an electron beam that melts the powder. Similar to SLM the EBM parts are fabricated layer by layer but in a high vacuum chamber, which makes this method suitable for highly reactive materials, such as titanium (Ti) and alloys. Another advantage of EBM over other AM techniques is that the process occurs at temperatures as high as 1000°C resulting in low residual stresses during fabrication and eliminating the need of post heat treatment. A comparison between capabilities and resulting properties of biomedical components fabricated via different AM techniques, explained above, is presented in Table 12.1 Also, Table 12.2 provides a summary of the laser type, fabrication resolution, and the type of materials that can be used in each of the discussed AM techniques.
12.3
Advantages, limitations, and challenges of AM
Similar to any other fabrication technique, there are advantages and challenges associated with using AM in medical devices. In this sections some of the main advantages and challenges associated with these technologies are discussed, which may help in making a decision in using AM for certain medical applications.
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Table 12.1 Comparison between the various AM techniques for fabricating biomedical components (Trevisan et al., 2017). Method
Component
Mechanical properties
Ref.
SLM
Porous Ti-6Al-4V scaffolds Porous Ti-6Al-4V structures Porous Ti structures
E ¼ 3.7–6.7 Pa UCS ¼ 145–164 MPa E ¼ 7–60 GPa YS ¼ 471–809 MPa E ¼ 2.6–44 GPa YS ¼ 24–463 MPa E ¼ 3.9–12.9 GPa YS ¼ 49.6–107.5 MPa UCS ¼ 127.1–148.4 MPa E ¼ 14.5–38.5 GPa YS ¼ 138–194 MPa UCS ¼ 163–286 MPa
Wieding et al. (2012) Bandyopadhyay et al. (2010) Xue et al. (2007)
DED DED EBM
Compact and porous Ti-6Al-4V structures
EBM
Porous Ti-6Al-4V implant
Ponader et al. (2009) Li et al. (2012)
E, modulus of elasticity; UCS, ultimate compressive stress; YS, yield stress.
Table 12.2 Common materials, resolution, and laser type for each of the mentioned powderbased metal AM methods (Shayesteh Moghaddam et al., 2017). Method
Common materials
Resolution
Laser type
SLM
Stainless steels, aluminum (Al), copper (Cu), iron (Fe), cobalt (Co)chrome (CR), Ti, Ni-based alloy, and a mixture of different types of particles (Fe, Ni, Cu, and Fe3P) (Abe et al., 2001; Childs et al., 2005; Kruth et al., 2004) Steel alloys, stainless steel, tool steel, Al, bronze, Co-Cr, (Fe, Ni)TiC composites and Ti (Xiao and Zhang, 2007) Ti, Ti-6Al-4V, Ti Grade 2, Co-Cr
20 μm (Ravari et al., 2015)
Nd:YAG (Delgado et al., 2012)
20 μm (H€anninen, 2001)
CO2 (Zhu et al., 2003), Fiber (Schleifenbaum et al., 2010) N/A
DED
EBM
N/A
12.3.1 Advantages of AM The advantages of AM in medical device industry originate from both the needs of the medical industry—such as complex geometries, patient-to-patient variation, high required precision, etc.—as well as the capabilities of AM, such as flexibility in geometry and fabrication of complex geometries, such as porous lattice structures, that conventional methods are unable to make. Using AM, completely customized parts can be fabricated that can be very advantageous for medical devices, where the device may vary from patient to patient. Moreover, the use of different materials to fabricate medical devices and parts is easily possible in AM, as long as the material can be produced
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to a form that can be used by AM machines. For an optimized AM process, very little postfabrication treatment is needed. For AM, a 3D model of the final product is generated in a computer before fabrication. This provides the opportunity of virtually assessing the appropriateness of the product and its match with the final implantation with the anatomy of the patient. In such cases, modifications can be made to the product quickly, before making the device. Finally, because of all these characteristics, AM of medical devices has a noticeably shorter supply-chain lead time and, in general, lower production cost.
12.3.2 Challenges of AM Although AM technologies are advantageous in making medical devices, there are challenges toward achieving a successful fabrication with desired functional and structural properties. Indeed, first time fabrication of parts with a material could be challenging until a set of appropriate processing parameters (e.g., laser power, scanning speed, etc., in a laser-based manufacturing) could be determined. This procedure adds to the lead time and cost of fabrication in an AM process. On the other hand, since AM is in its early stages of development, there are a lot of uncertainties on the effects of various factors, such as shape and size of the part, on the properties of the fabricated parts. Understanding and tailoring the influence of various fabrication parameters on the mechanical properties of the part requires extensive research. This also requires regulations for fabricating and characterizing AM parts to ensure consistency of fabrications using different machines and at different fabrication sites. Dimensional accuracy of the AM parts is another challenge with fabricating metallic medical devices especially for miniature components. In addition, providing an acceptable surface quality of the fabricated part, to eliminate postfabrication modifications, requires a lot of work and may not be possible for most materials and geometries at the current state of metal AM. Finally, durability of the fabricated AM parts is still the main challenge toward using these parts for medical applications, as majority of the real-world applications involve cyclic loading applied to these components. At the current state of AM, imperfections such as voids and impurities are inevitable in a fabricated AM component. Moreover, surface finish of the as-fabricated AM parts shows rough features. All these fabrication imperfections contribute to the fatigue failure of the AM devices and reduce their durability (Bagheri et al., 2018; Mahtabi et al., 2015a; Yadollahi et al., 2018).
12.4
Applications of AM in medical devices
The applications of AM in various industries has grown rapidly during the past decade. Especially in medical industry, these applications range from educational aspects to fabrication of final medical devices and implants ( Javaid and Haleem, 2017; Tuomi et al., 2014; Zadpoor and Malda, 2017). While applications of AM in biomedical applications can be categorized based on different perspectives, some of the major application of AM in medical devices are discussed in this section.
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Fig. 12.5 A prototype forceps fabricated using additive manufacturing (Kontio et al., 2012).
12.4.1 Tools, guides, and instruments for medical devices One of the applications of AM technology in medical devices is to fabricate tools and instruments to improve the effectiveness of a medical procedure or surgery. The main advantage of AM for this type of application is that the complete process from design to fabrication can be done at once and the production process is much faster and more consistent. AM drill guides (Bibb et al., 2009; Labadie et al., 2008), oral appliances (Salmi et al., 2013), and surgical instruments/guides (Kontio et al., 2012) are examples of such applications. Fig. 12.5 shows an example of the additively manufactured tools and instruments. Using AM, instruments can be made for a specific operation—which are not available in the market—to meet the especial needs of the operation. As the devices and instruments for this type of application are in direct contact with body fluids and tissues for a limited time, one important consideration is that the material used for this type of application must be sterilizable, but does not need to be implantable (Tuomi et al., 2014).
12.4.2 Prosthesis and orthotics The AM has been also used in orthotic and prosthetic applications. Similar to other applications, these applications benefit from the capability of AM in creating customized parts. Moreover, for prosthetic devices, AM enables fabricating functionally tuned devices by matching the dimensions of the device. For such applications, multimaterial technologies that are able to print colorful materials, which are not necessarily metals, are very advantageous. An additively manufactured prosthetic hand is shown in Fig. 12.6.
12.4.3 Customized implants Freedom in form is the main advantage of AM for fabricating customized, patientspecific implants. Using this capability, implants can be fabricated to match the anatomy of a patient. Examples of implants that are fabricated using various AM techniques are presented in Fig. 12.7 Implant applications of AM are so advanced that
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Fig. 12.6 Examples of AM orthotic and prosthetic parts: (A) Prosthetic hand (Burn et al., 2016), (B) foot orthosis (left) and ankle-foot orthosis (right) (Chen et al., 2016).
there are already companies that offer patient-specific AM implants following the procedure explained in Section 12.2. In addition, complex geometries such as lattice structures with intentional porous structures can be fabricated using different AM techniques. Examples of such lattice structures are shown in Fig. 12.7B. The lattice structures can be designed to meet different requirements, such as matching the mechanical properties of the implant to the biological tissue, and preventing some postimplantation issues, such as stress shielding. The porous structure of the implants also provides the potential for growth of tissue during healing, as well as osseointegration. The AM was used to fabricate artificial teeth as well (Liang et al., 2018) (Fig. 12.7C).
12.4.4 Education and training Another important use of AM is its use for educational and training purposes, as well as for communicating with patients. In addition, 3D reconstruction of the devices and physical parts (Zadpoor and Malda, 2017; Spottiswoode et al., 2013; Zein Nizar et al., 2013) using an AM can be very helpful for planning complex surgeries. The
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(A)
(C)
(B)
(D)
Fig. 12.7 Examples of medical implants fabricated by AM techniques: (A) Ti alloy dental implant (Obaton et al., 2017), (B) NiTi bone fixation plate, (C) lattice structure of NiTi fabricated by SLM technique, and (D) as-fabricated titanium sternum/rib implant made by EBM technique (Thompson, 2018).
capabilities of AM can also help evaluate the implantation procedure and the match between patient’s anatomy and the medical device. The AM parts can be fabricated to represent human anatomy and for educating new surgeons and medical students by providing a more realistic environment. Multimaterial printing can be very helpful in this type of application by providing colorful fabricated objects.
12.4.5 Drug delivery In 2015, food and drug administration (FDA) approved drug products fabricated by AM (Goyanes et al., 2016), indicating the great potential of AM for drug delivery devices in the solid form (Zadpoor and Malda, 2017). The main AM techniques that can be used for this type of application are binder jetting techniques, material extrusion, and material jetting. Using an AM, the drug release profile can be controlled by
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Fig. 12.8 Revenue split of AM equipment customers (Europe, 2015).
adjusting the shape of the device. Moreover, spatial distribution of the active agent in the device can also be designed in AM devices to have more control over the release profile. Another advantage of AM for drug delivery devices is the capability of fabricating multidrug materials into a single delivery device.
12.5
Market and statistics
The AM market is one of the world’s fastest growing markets with a predicted market value of $10.8 billion by 2021. Fig. 12.8 shows the revenue of AM medical devices (Europe, 2015). It can be seen that the most rapid growth is in automotive, electronics, and medical industries. Furthermore, considering very broad and generic applications as well as increasing number of accidents causing injuries and diseases there are demands in the market for faster AM of medical devices. The advantage of designing patient-specific devices is,
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another reason toward additively manufactured medical devices in the market, to such an extent that it is estimated that in the near future AM will be used for 80% of the implants. Reports show that a total market of approximately $1.5 billion is anticipated for AM of medical devices by 2026 (Liang et al., 2018).
12.6
Biological considerations and biocompatibility of AM parts
12.6.1 Biological consideration Technological advancement and availability of AM methods have led to increased investment and use by medical device industry. In December 2017, the United States FDA issued a comprehensive guidance for industry and FDA staff on “Technical Considerations for Additive Manufactured Medical Devices” (Food and Drug Administration, 2017). Beside other technical and mechanical considerations, this FDA document introduces and discusses all the parameters and conditions for the AM process that can affect the biocompatibility of the fabricated parts. While different AM techniques have been used in medical device industry, according to FDA—in addition to biocompatibility—other critical considerations, including cleaning and sterility, must be evaluated for AM medical devices. In general, achieving adequate cleaning, sterility, and biocompatibility is a major concern for different methods of AM and can be a limiting factor for the fabrication of specific geometries and materials. For instance, porous or internally complex devices must be carefully designed and processed due to the challenging procedure of cleaning and sterilization of these parts. In this section first we discuss the general aspects of biological considerations of AM devices, including cleaning, sterility, and biocompatibility, and then as an example, the biological consideration of additively manufactured Ti-6Al-4V will be discussed in more detail. For all methods of AM fabrication, each of the process parameters including the raw material, building atmosphere, energy input properties, as well as any postfabrication treatments may significantly affect the biological properties of the final part. The first item to consider is the material composition of the final AM part after postfabrication treatments. Process parameters such as the energy input and the existing elements in the build chamber may cause unexpected chemical reactions, which may adversely affect the biocompatibility of the fabricated parts. If this is the case, the content and build atmosphere properties (i.e., temperature, humidity, etc.) of the AM process must be controlled and monitored during the fabrication. In some cases, an inert gas such as argon is fed into the built chamber to decrease the oxygen level to specific levels to prevent the oxidation of the main material. Postfabrication processing, especially heat treatment, may have the same effect and cause the formation of a nonbiocompatible element. Although AM facilitates the creation of devices with complex geometries, such as components with engineered porosity, internal voids or cavities that may not exist in traditional manufacturing methods, are currently inevitable in AM products.
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Therefore, one should also consider the increased difficulty in removing manufacturing material residues (cleaning) and in the sterilization due to the likelihood of increased surface area, generation of extensive tortuous pathways, and creation of internal voids with limited or no access. In cases where the residual materials could have an adverse effect on the biological aspect of the fabricated part, the manufacturer must establish and maintain procedures for the use and removal of such residual manufacturing material to ensure that it is removed or limited to an amount that does affect the part’s features. For complex AM fabricated geometries, it should be mentioned that since many end-user facilities may not have access to equipment or materials required to implement specific cleaning procedures, only devices that are sufficiently clean of residual manufacturing materials should be provided to the end user. To evaluate the biocompatibility of the final finished devices, the same procedure as in conventional devices must be considered. For any medical device, FDA suggests to evaluate the biocompatibility as described in the guidance “Use of International Standard ISO-10993, ‘Biological Evaluation of Medical Devices Part 1: Evaluation and Testing within a Risk Management Process’” (Anand, 2000). Based on the FDA document, the first step for the assessment of the biocompatibility is the risk management for biocompatibility evaluations. One should identify and asset the potential risks based on the application of the medical device. For instance, a medical device that is in temporary contact with skin needs different types of evaluations compared to a permanent cardiovascular implant. After the risk assessment, a series of required tests should be designed and reported. These tests could be one or a couple of biocompatibility tests including, cytotoxicity, sensitization, hemocompatibility, pyrogenicity, implantation, genotoxicity, carcinogenicity, reproductive and developmental toxicity, and degradation assessments.
12.6.2 Biocompatibility of AM Ti and its alloys Ti and its alloys are among the most well-known biocompatible metals that can be fabricated via AM. In general, in addition to the biocompatibility, bioadhesion, biofunctionality, and corrosion resistance are the other important properties for biomedical applications. The biocompatibility of Ti is mainly due to its low electrical conductivity, which contributes to the electrochemical oxidation of Ti leading to the formation of a thin passive oxide layer on its surface (Quinn and Armstrong, 1978). This oxide layer creates a high resistance to corrosion. In addition, this protective oxide layer remains at pH values of the human body (Schiff et al., 2002). In this section, as an example, some of the studies on biocompatibility of AM Ti and its alloys are discussed in more details. Commercially pure Ti (CP-Ti) and Ti-6Al-4V are the most well-known compositions for the fabrication of Ti implants. The American Society for Testing and Materials (ASTM F67-06, 2006) specifies CP-Ti and Ti-6Al-4V as grade 1–5. Grades 1–4 are unalloyed CP-Ti and grade 5 is the Ti-6Al-4V alloy (ASTM F136-08, 2008). Table 12.3 summarizes the mechanical properties and different grades of Ti based on ASTM standards F67 and F136 (ASTM F67-06, 2006, 2008). The first step to
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Table 12.3 Selected mechanical requirements properties for an implantable Ti Bar (Elias et al., 2008).
Material
Specification
CP-Ti
ASTM F67 Grade 1 ASTM F67 Grade 2 ASTM F67 Grade 3 ASTM F67 Grade 4 ASTM F67 Grade 5
CP-Ti CP-Ti CP-Ti Ti-6Al4V
Tensile strength (MPa)
0.2% proof stress (MPa)
Elongation (%)
Elastic modulus (GPa)
240
170
24
103–107
345
275
20
109–107
450
380
18
103–107
550
483
15
103–107
860
795
10
114–120
fabricate a biocompatible Ti is to reach to the standard mechanical properties based on the standards. In addition to the mechanical properties, the chemical composition of the fabricated part must fulfill some standard requirements. ASTM and ISO have recently published new standards (ISO 22068-14, 2014; ASTM F3001-14, 2014; ASTM F2924-14, 2014; ASTM 2885-11, 2011) for additively layered manufactured (ALM) Ti parts, which are summarized in Table 12.4. These standards also include the minimum mechanical properties of the fabricated parts via ALM and is summarized in Table 12.5. Following the mechanical properties and chemical composition, in vitro and in vivo studies have been performed on the AM fabricated Ti parts that verify the biocompatibility and reliability of this material. These studies used cultured fibroblasts Table 12.4 Chemical composition requirements for ALM Ti according to ASTM standards. Chemical composition (wt%) Specification
Al
V
Fe
C
N
O
H
Y
ASTM F2885 Grade 5 ASTM F2924 ALM Ti-6Al4V ASTM F3001 ALM Ti-6Al4V
5.5–6.75
3.5–4.5
<0.30
<0.08
<0.05
<0.20
<0.015
<0.005
5.5–6.5
3.5–4.5
<0.3
<0.08
<0.05
<0.20
<0.015
–
5.5–6.5
3.5–4.5
<0.25
<0.08
<0.05
<0.13
<0.012
<0.005
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Table 12.5 Mechanical properties of ALM Ti-6Al-4V.
Material
Specification
Ti-6Al4V ALM Ti-6Al4V ALM Ti-6Al4V
ASTM F2885 Grade 5 ASTM F2924 ALM Ti-6Al4V ASTM F3001 ALM Ti-6Al4V (ELI) Tensile
Tensile strength
0.2% proof stress
Elongation (%)
Reduction of area
780
680
10
15
895
825
10
15
860
795
8
25
and osteoblasts in the observation of cell responses to surfaces and also human and animal subjects. Table 12.6 summarizes the processing methods, test types, and general findings of these studies.
12.7
Case study: Stiffness-matched bone fixation plates for mandibular reconstruction surgery
Mandibular segmental defects usually happen due to tumors, cancers, radiation therapy, and trauma. The standard of care for the treatment of a segmental defect starts with resectioning the defected region and using a bone graft (i.e., a healthy piece of bone harvested from patient’s own body, such as fibula) to fill the resected region. As a part of this procedure, mechanical fixations including metallic bone plates and screws are used to immobilize the pieces of bone, and to facilitate and expedite the healing process (Fig. 12.9A) (Shayesteh Moghaddam et al., 2016). Conventionally, these bone plates are made of Ti-6Al-4V alloys. Although biocompatible Ti-6Al4V bone plates provide enough immobilization immediately after the surgery, they may cause major complexities when the bones heal. Due to the high stiffness of Ti-6Al-4V in comparison to cortical bones, the load and stress distribution on the surrounding bones is affected in a way that a major portion of the loads are carried by the fixation plates instead of the bones. This reduced level of stress, which is also called stress shielding, leads to bone resorption and eventually failure of the surgery based on the bone remodeling phenomena (Mertens et al., 2014). The existence of bone resorption due to stress shielding can be clearly seen in the hip implant, thanks to X-ray images, in Fig. 12.9B. As it can be seen in this figure, the surrounding bones, especially on the top right portion of the hip implant, are darker in comparison with the healthy case that shows the bone resorption due to the stress-shielding effect. One solution to prevent this failure is to remove the bone plates after the bonehealing period in a secondary surgery. Although removing the bone plates could
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Table 12.6 Summary of AM processing methods and findings. Method
Alloy
Biocompatibility test
Cell line/ implantation
EBM
Ti-6Al-4V
In vitro
Human fetal osteoblasts (hFOB 1.19)
EBM
Ti-6Al-4V
In vivo
Frontal skull of domestic pig
EBM
Ti-6Al-4V
In vivo
Rabbit femur and tibia
EBM
Ti-6Al-4V
In vitro
Human adiposederived adult stem cells (hASC)
EBM
Ti-6Al-4V
In vitro and in vivo
EBM
Ti-6Al-4V
In vivo
Osteoblasts extracted from Calvaria of rabbits, Calvaria of rabbits Sheep
DMLS
Ti-6Al-4V
In vitro
Human osteoblasts cells (HOB)
DMLS
Unspecified
In vitro and in vivo
DMLS
Ti-6Al-4V
In vivo
BMP-7 transduced human gingival Fibroblasts Human anterior mandible, minipig mandibular
General findings Reduced cell proliferation in highly rough surfaces (ASTM F136-13, 2013) More bone contact in more porous samples (Ponader et al., 2009) As-EBM implant response comparable to machined (Thomsen et al., 2008) Increased proliferation on porous compared to polished and unpolished EBM discs (Haslauer et al., 2010) Proliferation in porous EBM Ti-6Al4V implants matched coated implants (Li et al., 2012) High bone-implant contact in porous implant (Palmquist et al., 2011) Cultured cells attached and proliferated on SLM substrates (Hollander et al., 2003) In vitro and in vivo test data showing substantial bone ingrowth (Hollister et al., 2005) Peri-implant bone in close contact with the surface of the implant (Mangano et al., 2010)
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Table 12.6 Continued Method
Alloy
Biocompatibility test
Cell line/ implantation
DMLS
Ti-6Al-4V
In vitro
Human osteoblasts human
LENS
CP-Ti
In vitro
LENS
Ti-6Al-4V
In vivo
Human osteoblast cells (OPC1) Male SpragueDawley rats
Modified FDM
CP-Ti
In vitro
L929 mouse fibroblast
General findings Osteoblasts wellspread and with multiple contact points (Warnke et al., 2008) Cells well spread on porous Ti (Xue et al., 2007) Increase in calcium (bone) within implant pores (Bandyopadhyay et al., 2010) Excellent bone cell attachment and proliferation (Wiria et al., 2010)
Fig. 12.9 (A) Schematic illustration of a mandibular reconstruction ( Jahadakbar et al., 2016) and (B) the effect of stress shielding and bone resorption (Engh et al., 1987).
prevent the risk of failure, extra cost and risk of the second surgery, and the increased recovery time are the main side effects of this solution (Boyd et al., 2012). Recently, patient-specific, stiffness-matched bone plates have been introduced to prevent the stress-shielding phenomenon and to eliminate the need for a second surgery ( Jahadakbar et al., 2016). These bone plates are made of Ni-Ti alloy (NiTi or Nitinol), designed based on medical images of the patient (CT scans or X-rays), and are fabricated via AM. The stiffness-matching option minimizes the stress shielding on the
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Fig. 12.10 The process of design and fabrication of patient-specific, stiffness-matched mandibular bone fixation plates.
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surrounding bone and prevent bone resorption. Therefore, there is no need to remove the bone plates with a second surgery. It is worth noting that NiTi, which is the base material for this next generation of bone plates, is a biocompatible shape memory alloy with unique features such as superelasticity and high fatigue life at large levels of strains, which make it a great candidate for biomedical implants (Mahtabi et al., 2015b,c; Mahtabi and Shamsaei, 2017). Due to the high thermomechanical sensitivity of this alloy, the conventional methods of fabrication such as machining and forging are very challenging for this material (Elahinia et al., 2012). Due to these manufacturing limitations, in the past NiTi was only available in the form of wire, rod, tube, and sheet. Recently, AM has been used for successful fabrication of this alloy with almost no limitations directly from CAD files (Elahinia et al., 2016). The mechanical properties and biological properties of the AM fabricated NiTi have been compared with conventionally fabricated NiTi and confirmed the successful fabrication of this alloy via SLM technique (Habijan et al., 2013; Haberland et al., 2012; Ibrahim et al., 2018). Fig. 12.10 shows a general schematic of the process of AM of mandibular fixation plate. The design of these bone plates starts with CT-scan data of the patient’s mandible. Based on the CT-scan data, first a 3D CAD model of the patient’s mandible is created. Then in a virtual process the defected region is resected and replaced with a double barrel fibula graft. Following this step, the general geometry of the required bone plates is designed. Then the reconstructed mandible along with the bone plates and screws are transferred to a finite element (FE) software and modeled. In a series of simulations, using the CT-scan data, the required stiffness of the bone plates is calculated. Next, specific levels and types of porosity are assigned to the bone plates to modulate the final stiffness of the bone plates to the required level. Finally the CAD file of the patient-specific, stiffness-matched bone plates are designed based the patient’s anatomy and FE simulations is used for AM ( Jahadakbar et al., 2016).
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