Additive manufacturing of metallic and polymeric load-bearing biomaterials using laser powder bed fusion: A review

Abstract

Surgical prostheses and implants used in hard-tissue engineering should satisfy all the clinical, mechanical, manufacturing, and economic requirements in order to be used for load-bearing applications. Metals, and to a lesser extent, polymers are promising materials that have long been used as load-bearing biomaterials. With the rapid development of additive manufacturing (AM) technology, metallic and polymeric implants with complex structures that were once impractical to manufacture using traditional processing methods can now easily be made by AM. This technology has emerged over the past four decades as a rapid and cost-effective fabrication method for geometrically complex implants with high levels of accuracy and precision. The ability to design and fabricate patient-specific, customized structural biomaterials has made AM a subject of great interest in both research and clinical settings. Among different AM methods, laser powder bed fusion (L–PBF) is emerging as the most popular and reliable AM method for producing load-bearing biomaterials. This layer-by-layer process uses a high-energy laser beam to sinter or melt powders into a part patterned by a computer-aided design (CAD) model. The most important load-bearing applications of L–PBF-manufactured biomaterials include orthopedic, traumatological, craniofacial, maxillofacial, and dental applications. The unequalled design freedom of AM technology, and L–PBF in particular, also allows fabrication of complex and customized metallic and polymeric scaffolds by altering the topology and controlling the macro-porosity of the implant. This article gives an overview of the L–PBF method for the fabrication of load-bearing metallic and polymeric biomaterials.

Full text

Journal of Materials Science & Technology 94 (2021) 196–215
Contents lists available at ScienceDirect
Journal of Materials Science & Technology
journal homepage: www.elsevier.com/locate/jmst
Additive manufacturing of metallic and polymeric load-bearing
biomaterials using laser powder b e d fusion: A review
Alireza Nouri
a , b , ∗, Anahita Rohani Shirvan
c
, Yuncang Li
a
, Cuie Wen
a , ∗
a
School of Engineering, RMIT University, Melbourne, Victoria 3001, Australia
b
Biomedical Engineering Department, Amirkabir University of Technology, Tehran, Iran
c
Text il e Engineering Department, Amirkabir University of Technology, Tehran , Iran
a r t i c l e i n f o
Article history:
Received 18 January 2021
Revised 12 March 2021
Accepted 13 March 2021
Available online 20 May 2021
Keywo rds:
Additive manufacturing
Load-bearing biomaterials
Powder bed fusion (PBF)
Selective laser melting (SLM)
Selective laser sintering (SLS)
a b s t r a c t
Surgical prostheses and implants used in hard-tissue engineering should satisfy all the clinical, mechani-
cal, manufacturing, and economic requirements in order to be used for load-bearing applications. Metals,
and to a lesser extent, polymers are promising materials that have long been used as load-bearing bioma-
terials. With the rapid development of additive manufacturing (AM) technology, metallic and polymeric
implants with complex structures that were once impractical to manufacture using traditional process-
ing methods can now easily be made by AM. This technology has emerged over the past four decades
as a rapid and cost-effective fabrication method for geometrically complex implants with high levels of
accuracy and precision. The ability to design and fabricate patient-specific, customized structural bioma-
terials has made AM a subject of great interest in both research and clinical settings. Among different
AM methods, laser powder bed fusion (L–PBF) is emerging as the most popular and reliable AM method
for producing load-bearing biomaterials. This layer-by-layer process uses a high-energy laser beam to
sinter or melt powders into a part patterned by a computer-aided design (CAD) model. The most impor-
tant load-bearing applications of L–PBF-manufactured biomaterials include orthopedic, traumatological,
craniofacial, maxillofacial, and dental applications. The unequalled design freedom of AM technology, and
L–PBF in particular, also allows fabrication of complex and customized metallic and polymeric scaffolds
by altering the topology and controlling the macro-porosity of the implant. This article gives an overview
of the L–PBF method for the fabrication of load-bearing metallic and polymeric biomaterials.
©2021 Published by Elsevier Ltd on behalf of Chinese Society for Metals.
1. Introduction
The term “additive manufacturing” (AM) is generally used to
explain a manufacturing technology that builds three-dimensional
(3D) objects through layer-by-layer addition of material. AM is the
most appropriate definition and the most correct term for what
used to be called rapid prototyping (RP) and what is now popu-
larly called 3D printing [1] . Chuck Hull was the first person who
mentioned the use of AM, in the early 1980s. He coined the term
“stereolithography” and patented the technology in 1986 [2] . Al-
though it has been four decades since the advent of AM, it has only
become progressively widespread within both surgical and non-
surgical fields in recent years [3] . Reportedly, the AM with metal
powders market was valued 800 million USD in 2018 and is pro-
∗Corresponding authors.
E-mail addresses: alireza.nouri@gmail.com (A. Nouri), cuie.wen@rmit.edu.au (C.
Wen).
jected to reach 5650 million USD by 2025, at a CAGR of 27.7% dur-
ing the forcast period [4] .
All AM technologies are based on the principle of slicing a
solid model in multiple layers, uploading the data to the machine,
and building the part via layer-by-layer addition of material fol-
lowing the sliced model data using a heat source (laser, electron
beam, or electric arc) and feedstock (metal powder or wire) [5] .
This makes AM a robust manufacturing technology for produc-
ing geometrically complex objects that would be extremely diffi-
cult or even impossible to fabricate through conventional manufac-
turing techniques. Conventional manufacturing requires formative
(molds) or subtractive (machining) techniques that are performed
in multiple steps using costly equipment and tooling. Furthermore,
these conventional manufacturing techniques do not allow for the
complex geometries and precise details that are often required in
biomedical engineering applications [6–8] . On the other hand, AM
has emerged as a promising fabrication technique that offers re-
duced operator error and increased repeatability, speed, and cost-
effectiveness over conventional techniques [9] . There are still other
https://doi.org/10.1016/j.jmst.2021.03.058
1005-0302/© 2021 Published by Elsevier Ltd on behalf of Chinese Society for Metals.

A. Nouri, A. Rohani Shirvan, Y. Li et al. Journal of Materials Science & Technology 94 (2021) 196–215
advantages of the AM technique including freedom of design, po-
tential elimination of tooling and production steps, and part con-
solidation by reducing assembly requirements.
The medical devices industry is a major sector that commands
a large share in the AM. The role of AM in healthcare is becoming
increasingly important due to its wide utilization in therapeutic
delivery [5–7] , surgical planning [8] , implant design [9] , and tis-
sue engineering [10–12] . Owing to the digital nature of the built
parts in AM, this technology is likely to produce medical devices
with predictable clinical outcomes that are much closer to manu-
facturers’ specifications; what is known as customization. The im-
portance of customization in the medical device sector is primarily
because each patient has their own individual anatomy as com-
pared to others. Most implants are currently designed and fabri-
cated based on an average bone shape and thus, do not necessar-
ily fit all patients [10] . As a consequence, mass-produced implants
usually lead to significant compromises in the anatomical accuracy
and recovery time. Customization allows the production of a wide
range of implants in a timely and a cost-effective manner with
specific properties and shapes that meet the patient needs [11] .
Studies and reports from the medical devices industry show
that 11% of their revenue comes from 3D-printed implants and
medical devices. As discussed earlier, this growing interest is as-
sociated with the need to produce patient-specific customized bio-
materials today. AM also fits well into tissue engineering, allow-
ing for the fabrication of porous and solid implants that promote
bone regeneration, and complex 3D bioprinting of cells [7] . Among
these approaches, the emergence of bioprinting in recent years has
been realized as a new alternative to fabricate tissues. Here, com-
plex living structures are fabricated through layer-by-layer deposi-
tion of living cells [12] . While we are still years away from being
able to produce viable additively manufactured organs, it has been
many years since hard-tissue engineering has whole-heartedly em-
braced the use of digital manufacturing technologies. As people
age, their load-bearing joints become more susceptible to wear,
fractures, and arthritis. This is a driving force which demands the
design and fabrication of new implants with suitable properties
for specific applications. These implants should mimic the struc-
ture and mechanical properties of natural bone so as to aid and
expedite the healing process [ 13 , 14 ].
The load-bearing biomaterials used in hard-tissue engineering
are largely focused in the areas of osteo and dental clinical ap-
plications, and are primarily used as: (i) load-bearing prostheses
to replace injured, diseased, or abnormal body parts such as to-
tal joint replacements and craniofacial plates; and (ii) fixation de-
vices that are used to stabilize broken bones while healing, such
as bone plates, intramedullary nails and staples [15] . Materials that
are used for the fabrication of a load-bearing implant play a vital
role in its success or failure in the human body. From a mechani-
cal standpoint, these materials should possess high wear and cor-
rosion resistance, high fracture toughness and fatigue strength, and
low elastic modulus. Moreover, they must be made of highly bio-
compatible alloying elements, as the lack of any or all of these fac-
tors may lead to incompatibility between the bone tissue and im-
plant materials [16] . In other words, these biomaterials should sat-
isfy all the clinical, mechanical, manufacturing, and economic re-
quirements in order to be used for load-bearing applications. With
the rapid development of AM technology, metallic and polymeric
biomaterials with complex and precise structures that were once
impossible to manufacture via traditional processing methods can
now be made easily. The resulting personalized treatment for pa-
tients ensures the implant has a better bone repair effect [17] .
For load-bearing biomaterials, the transfer of force should oc-
cur in a physiological way. Thus, the bone’s integrity needs to be
respected, especially on the interface between the implant and the
bone. The majority of defects are located on this interface because
of the sudden change in physiological stiffness of the facing sub-
strates. Simulations of mechanical stress at the bone-implant in-
terface confirm the effect of stress-shielding [18] .
Metals, and to a lesser extent polymers and ceramics, are key
families of implant biomaterials that are widely used in many
load-bearing applications. By virtue of their interatomic bonds,
metals are characterized by a combination of high strength and
moderate plastic deformation that favors their application for load-
bearing structures. Thus, it is accurate to state that metallic mate-
rials are the most common and useful load-bearing biomaterials
[19] . Alongside metallic biomaterials, polymeric biomaterials are
also utilized for load-bearing applications. Some structural poly-
meric biomaterials have gained much interest in the field as they
withstand high physiological loads without deformation or fractur-
ing [ 20 , 21 ].
The concept of using AM technology in the fabrication of or-
thopedic and dental implants has been discussed in the literature
over the last few years; however, no fundamental study has ever
been collectively undertaken to address this technology in rela-
tion to load-bearing biomaterials. The current review not only de-
scribes the latest updates and applications of AM technology in or-
thopedics and dentistry, but also covers the traumatological, cran-
iomaxillofacial, and structural porous scaffolds load-bearing appli-
cations in hard-tissue engineering. An attempt has also been made
to deliver the theoretical and practical aspects of AM technology in
manufacturing of both polymeric and metallic load-bearing bioma-
terials through a straightforward approach.
2. Powder bed fusion
The most common AM technology applied to the production of
load-bearing biomaterials is powder bed fusion (PBF) [17] . Gener-
ally, in laser-based techniques , the layer thickness on the plat-
form is typically 100 μm for polymer powders and 20-100 μm for
metal and ceramic powders. However, in electron beam-based PBF
(i.e. EBM), a layer thickness of 50-200 μm is typically used in the
process [22] . It is worth noting that for directed energy deposi-
tion (DED) technique, as another metal additive manufacturing sys-
tem, powders are larger and thicker layer thickness are being ap-
plied. Han et al. [23] repoerted a typical layer thickness of 10 0-80 0
μm for the DED of different high-entropy alloys. The build plat-
form is then lowered and a new layer of powder is spread across
the previous layer using a roller. In each layer, the powders are
melted or sintered and bind together as they cool. The process is
repeated until the entire part/object is finished [ 5 , 11 ]. Following
the completion of laser scanning, loose powders are removed from
the build platform and the component is taken from the substrate
plate manually or by electrical discharge machining (EDM) [24] .
PBF technologies include selective laser melting (SLM), selective
laser sintering (SLS), direct metal laser sintering (DMLS), and elec-
tron beam melting (EBM). In all these technologies, heat is used to
fuse the powdered materials. The differences between these tech-
nologies lie in their energy source and their powder materials [25] .
Unlike most other AM processes, here the powder bed around the
built parts serves as a support structure, enabling the creation of
support-free parts. As such, the material cost of the support struc-
ture and the post-processing cost of removing supports are both
eliminated. Apart from the economic standpoint, the elimination
of support structures brings the geometric freedom of design and
faster production of the parts [26] .
SLS is mainly used in the processing of polymers and ceramics,
whereas SLM, DMLS, and EBM are specifically used in the man-
ufacturing of metals and alloys [27] . In SLM, SLS, and DMLS, a
laser beam is used to fuse powder particles, whereas EBM uses an
electron beam as the energy source [ 25 , 28 ]. Laser-based systems
are operated under an inert atmosphere (especially for titanium
197

A. Nouri, A. Rohani Shirvan, Y. Li et al. Journal of Materials Science & Technology 94 (2021) 196–215
Fig. 1. Schematic setup of L–PBF additive manufacturing technique.
(Ti) processing), as opposed to the vacuum environment of the
electron-beam systems. Although the vacuum systems in the EBM
process are more expensive, they show lower residual stress than
laser-based systems. As a result, electron beam–processed parts
can be used without any stress-relieving operation [ 5 , 27 ].
In the sintering processes (i.e. SLS and DMLS), powder materi-
als are consolidated on top of each other by solid-state sintering,
liquid-phase sintering (LPS), or partial-melting mechanisms [29] .
This leads to internal porous structures and rough surfaces. In the
melting processes (SLM and EBM), the powders are well fused
and create denser parts with enhanced mechanical properties [29] .
These powder-based printing technologies are used for manufac-
turing of various materials including metals, polymers, ceramics,
and a combination of them for different applications [ 26 , 30 , 31 ].
Time efficiency, fast production, flexibility in design, capability of
manufacturing complex geometries, and capability of processing
all kind of materials are the main advantages of PBF technology
in comparison with its conventional counterpart methods such as
powder metallurgy [31] . Moreover, in this AM technology the re-
maining unprocessed powder can be reused [29] .
A schematic illustration of a laser-based PBF (L–PBF) system is
shown in Fig. 1 . A brief overview of various L–PBF technologies is
necessary in order to better understand their current and potential
applications in the manufacturing of load-bearing biomaterials.
2.1. SLS
The SLS process was developed by Carl Deckard for his master’s
thesis in the mid-1980s at the University of Texa s and patented
in 198 9 [32] . In SLS, parts are fabricated layer-by-layer by sinter-
ing the powder materials through a scanning laser beam. The laser
beam hits a thin layer of powder materials and raises the tempera-
ture of the powder to the melting point, resulting in particle bond-
ing, fusing the powders to each other and to the previous layer
to form a solid component. Then, the next layer is built on the
top of the sintered layer after an additional layer of powder is de-
posited via a roller mechanism [ 33 , 34 ]. In SLS, binding of the pow-
der particles may occur via several mechanisms such as viscous-
flow, particle wetting/infiltration and curvature effects [35] . In this
process, loose or slightly compacted powders with particle sizes in
the range of several microns to several hundred microns are typi-
cally used [36] .
A wide range of thermoplastic materials including engineering
and high-performance plastics can be processed via SLS. Various
metallic parts are also being produced by this technology [37] ;
however, very few studies have been carried out regarding the
use of SLS for metallic implants and equipment. Great efforts have
been made to process new technical thermoplastics like polyether
ether ketone (PEEK) for medical devices, although this requires
high temperatures, complex control, and a great deal of wastage
[1] . Although a wide range of amorphous and semi-crystalline
thermoplastic polymers have been experimentally tested for the
SLS process, commercial applications of SLS are still limited to
a small number of thermoplastic polymers including polyamide
(PA12 and PA11), polycarbonate (PC), polystyrene (PS), PEEK and
variants thereof [ 36 , 38 ]. The obtained laser-sintered products are
durable enough to be used in load-bearing applications where the
parts are subjected to mechanical loads [39] . There is no shape
constraint in this process, as no support is needed during manu-
facturing. The reason is that the excess powder in each layer helps
to support the part during the build [36] . Nevertheless, SLS has
some drawbacks including the porosity of the surface of the ob-
ject and the long manufacturing time due to the printer’s required
heating time and the 3D object’s cooling time [1] . Polymers used in
this process have high melting points (above autoclave sterilization
temperature) and excellent material properties for use in anatomi-
cal study models, cutting and drilling guides, and load-bearing ap-
plications [40] .
2.2. SLM
By virtue of powerful high-quality lasers, SLM was made com-
mercially available in 2005 by MCP Tooling Technologies (United
Kingdom), which is now known as MTT Technologies Group [41] .
The laser systems for SLM progressed from CO
2
laser ( λ≈10.6 μm)
to Nd:YAG fiber laser ( λ≈1.0 6 μm) and subsequently to Yb: YAG
fiber laser. This is due to the higher absorptance of metallic pow-
ders in relation to the radiation of such wavelengths in the infrared
region [24] .
The SLM process follows a similar principle to SLS; however,
in SLM powder is completely melted instead of partially melting
or sintering [42] . In other words, SLM is SLS conducted with a
high-energy laser beam with the objective of achieving complete
melting of powders [36] . The implementation of the idea of full
melting has been made possible by the continuously improved
laser-processing conditions in recent years, e.g. higher laser power,
smaller focused spot size, and thinner layer thickness [43] . The
thickness of the layer usually ranges from 20 to 100 μm. This is
chosen as a balance between achieving fine resolution and allow-
ing for good powder flowability. Furthermore, the energy of the
beam is much higher and the process is performed under a con-
trolled atmosphere [44] .
Generally, SLM can produce strong and dense parts, while SLS
is more suitable for processing of porous and weak objects. There-
fore, SLS is frequently used for manufacturing of polymers and
ceramics, while SLM is a common method for manufacturing of
metallic parts (e.g. 316L stainless steel, cobalt–chromium (Co–Cr),Ti
alloys, and commercially pure Ti (cp-Ti) [45] . However, SLM has
some limitations for the fabrication of ceramic materials. Ceram-
ics have limited flowability after melting and they are not able
to form a compact layer. In addition, the formation of porosity in
SLM-manufactured ceramics has a negative impact on the mechan-
ical properties of the component. Moreover, the balling effect and
cracking are other problems which are caused by the low conduc-
tivity of ceramic materials and huge temperature fluctuation dur-
ing the process, respectively [44] . Like EBM, the SLM requires that
the given materials exhibit similar properties with respect to laser
absorption and flow behavior of the liquid phase so as to obtain
198

A. Nouri, A. Rohani Shirvan, Y. Li et al. Journal of Materials Science & Technology 94 (2021) 196–215
the desired properties. Therefore, material availability in this pro-
cess is more limited than in the sintering process [46] . During the
SLM, the build chamber is often purged with an inert gas to pro-
vide a protective atmosphere for the heated metal parts against ox-
idation. Furthermore, some SLM systems are capable of preheating
the substrate plate or the entire build chamber [24] . The develop-
ment of SLM has been driven by the need to fabricate fully dense
parts with comparable mechanical properties to those of bulk ma-
terials and the need to avoid time-consuming post-processing. In
contrast to DMLS, the density of SLM-processed pure metals can be
increased up to 99.5% via the full melting mechanism. More impor-
tantly, one of the major advances of SLM lies in the possibility it
offers of processing pure metals, e.g. Ti, aluminum (Al), and copper
(Cu), which to date cannot be processed well with DMLS [43] .
SLS and SLM are high-resolution techniques and have reduced
manufacturing time. In other words, the whole part is fabricated
in one or two steps. SLS implants may exhibit a rough surface and
poor surface finish which can be desirable for some implants. In
contrast, SLM has a better surface finish than that of SLS. However,
there is a limitation in the materials which can be manufactured
by SLM [47] . In SLM, it is important to estimate the part’s den-
sity, accuracy, surface roughness, hardness, strength, and residual
stress. The crucial process parameters for the production of high-
quality and highly dimensionally accurate parts are the scanning
rate, spot size, laser power, layer thickness, and type of materials
[48] . In general, SLM is an attractive method for producing load-
bearing biomaterials with complex shapes due to its good mechan-
ical properties, high accuracy, and net shape ability [49] .
2.3. DMLS
DMLS is an extension of the SLS process and is conceptually the
same process, regenerating 3D parts from layer-by-layer additions
of fused metal powder. DMLS process was initially developed by
EOS GmbH of Munich (Germany) and it has been available com-
mercially since 1995 [50] . Under a PBF process, DMLS uses metal
powder and high laser power to sinter metal powders into the final
part. DMLS exclusively utilizes uncoated pre-alloyed metal pow-
ders as the sintering material, whereas in SLS polymers or coated
metal powders are used [ 46 , 51 ]. Maraging steel, stainless steel,
Al alloys, nickel (Ni) alloys, Co-Cr, Ti and Ti alloys are among the
different metallic materials in powder form that are used in the
DMLS process [52] .
In general, the DMLS mechanism is based on LPS, which is
based on a partial melting of the powder [43] . In this process, the
powder material is processed by an Ytterbium fiber laser or a CO
2
laser in an inert and thermally controlled chamber [ 50 , 52 ]. The
absorbed energy on the surface of the metal powder sinters the
underlying solidified powder, and in a few hours 3D parts with
high complexity and accuracy are produced by the DMLS machine.
By precise control of the porosity in each layer, DMLS can also
be used to fabricate porous metallic implants with a high degree
of pore interconnectivity, along with varying pore size, shape, and
distribution that mimic the architecture and mechanical properties
of natural bone [53] .
Heterogeneous microstructures and properties are regarded as
major problems with DMLS-fabricated parts. Moreover, support
structures are required and parts may need a variety of post-
processing treatments. These treatments may include removal
of support structures, surface finishing, heat treatment, post-
sintering, infiltration with a low-melting-point material or hot iso-
static pressing (HIP) to achieve the desired mechanical properties
and to eliminate porosity and anisotropy. Similar to SLS and SLM,
the DMLS process is also considered a suitable method for produc-
ing complex, fully functional metal parts with geometrical com-
plexity that would be difficult to produce by conventional meth-
ods. DMLS is the most preferred L–PBF technology with respect to
cost-effectiveness, short manufacturing time, and fabrication of a
wide variety of metal parts [ 43 , 54 ]. Table 1 summarizes the bene-
fits and drawbacks of the SLS, SLM and DMLS processes [55] .
The laser processing parameters, type of material and powder
characteristics are all important factors in L–PBF that significantly
affect the properties of the final products including mechanical
properties, surface roughness and dimensional accuracy. These pa-
rameters also have a direct influence on the manufacturing time
and quality of the final parts. In other words, any changes in the
process parameters such as laser power, hatching space, scanning
rate and scanning pattern can cause various undesirable effects like
balling effect, geometrical irregularity and insufficient density of
the final parts [56] . Fig. 2 illustrates the different types of L–PBF
induced-defects that are formed as the result of improper laser en-
ergy density.
These types of AM techniques provide the opportunity to design
and print patient-specific medical devices including load-bearing
implants from various structural biomaterials. The fabricated im-
plants are reproducible and have high quality in terms of struc-
ture and properties [57] . There are various load-bearing biomateri-
als that are designed and used for joint arthroplasty, bone-fracture
fixation, craniofacial and maxillofacial surgery, dental implants, and
structural porous scaffolds. The present article provides a compre-
hensive overview of the current status of load-bearing polymeric
and metallic biomaterials that are fabricated by L–PBF technol-
ogy. The review also highlights the main applications of L–PBF-
fabricated biomaterials for structural applications and investigates
their benefits and drawbacks as compared to conventional meth-
ods.
3. Materials in L–PBF
The materials selection for AM is still fairly limited. Although
this is partly due to market demand and costs, the available ma-
terials may not technically and perfectly meet the application re-
quirements. This limitation in materials selection does not ease or
simplify the decision-making process since the properties of mate-
rials are highly dependent on process and part geometry. In the L–
PBF processes, materials must exhibit a certain level of flowability
and packing efficiency to enable processing of unsintered powder
[39] . Sintering or melting of powdered material must be feasible in
the temperature range of the L-PBF machines. The material specifi-
cations presented by additive manufacturers only provide baseline
references and often requires additional critical observation during
applications. At present, processible materials in AM range across
all basic materials including metals, polymers, ceramics, and com-
posites. In principle there should be very little difference in mate-
rial compatibility among different PBF processes. However, some of
the processes are likely to support greater material selection partly
due to economic considerations [58] . This review focuses on poly-
meric and metallic materials used in L–PBF technology.
3.1. Polymers
In load-bearing applications, the number of materials that can
be used for AM of polymers is small compared to other ar-
eas. PEEK, ultrahigh molecular weight polyethylene (UHMWPE),
polypropylene (PP), polymethyl methacrylate (PMMA), polylactic
acid (PLA), polycaprolactone (PCL), polyvinyl alcohol (PVA), and
polyamide (PA) are the most common polymers which are pro-
cessed by SLS for load-bearing applications.
Molecular structure, crystallinity, and particle size are the most
important factors associated with polymer materials which have
direct effects on the final product fabricated by SLS. Molecular
structure is one of the most significant parameters that influence
199

A. Nouri, A. Rohani Shirvan, Y. Li et al. Journal of Materials Science & Technology 94 (2021) 196–215
Tabl e 1
A comparison between SLS, SLM and DMLS additive manufacturing methods.
Method Materials Advantages Disadvantages
SLS Polymers (e.g., PA12, PEEK, PCL, PMMA); Metals; Ceramics
•High dimensional accuracy
•Flexibility in the type of materials
•Easy modification and design
changes
•Elimination of any post-curing
•Best mechanical properties and
less anisotropy
•Rough surfaces
•Poor reusability of un-sintered
powder
•Shrinkage and warping due to the
thermal distortion
•The strength of the part is weaker
in the Z -direction than in the
other
•The price of machinery is
expensive
SLM Metals
•No distinct binder and melt
phases
•Producing fully dense parts in a
direct way
•Elimination of some
post-treatments
•Not suitable for well-controlled
composites
•Expensive laser and longer build
times
•Melt pool instabilities and higher
residual stress
DMLS Metals
•High speed
•Complex geometries
•High quality (high accuracy and
detailed resolution)
•Good static mechanical properties
•Medium surface roughness with
good biological performance
•Fully dense parts after heat
treatment
•High energy consumption
•Long build cycles
•Need of building support
•Need heat treatment to release
internal stress
•Grainy surface
Fig. 2. Different types of defects in L–PBF fabricated parts due to improper laser energy density (A: normal; B: poor material fusion; C: overheating; D: insufficient heating).
the melting viscosity and thus, indicates the suitability of a poly-
mer for SLS processing. Low zero viscosity ( η0
) and low surface
tension ( γ) of a polymer melt are necessary for a suitable coa-
lescence of polymer particles, which can equally affect the quality
of the final part. For example, a polymer with lower melt viscos-
ity provides a denser product. It is reported that the ideal melt
viscosity is 60 Pa.s from the Frenkel sintering model. For exam-
ple, polyamide (nylon) PA6 and PA12 are two polymers with simi-
lar structure but very different applicability in SLS processing. Be-
cause of the higher amount of hydrogen bonding in PA6, it exhibits
a higher melting point (223 °C in comparison to 187 °C for PA12)
and a higher melt viscosity. PA12 has different modes of packing,
which results in a narrower melting range, making it well suited
for SLS [39] . Regarding amorphous polymers processed with SLS,
the final product is usually brittle and instable because these types
of polymers have high viscosity above T
g and as a result a proper
coalescence does not take place [ 59 , 60 ]. Besides viscosity, as dis-
cussed earlier one should consider the marked effects of the size
and morphology of polymer powders on the quality and density
of the final part. The size and morphology of powder can influ-
200

A. Nour i, A. Ro hani Shirvan, Y. Li et al. Journal of Materials Science & Technology 94 (2021) 196–215
ence the powder flowability in L–PBF processes. For determination
of flowability, the Hausner ratio (H
R
) is used based on the follow-
ing equation [61] :
H
R
= ρtap
/ ρbulk (1)
where ρtap is the tapped density and ρbulk is the bulk density. Ac-
cording to the literatures, H
R
< 1.2 5 and H
R
> 1.4 indicate free-
flowing powder behavior and fluidization problems, respectively.
Amorphous polymer powders (e.g. PS, PMMA, PC) do not have a
sharp melting point but instead soften gradually with the increase
of temperature. The viscosity of amorphous polymers decreases
strongly with increasing temperature above the glass transition
temperature ( T
g
), but they are rarely as easy flowing as semi-
crystalline materials. Their lower flow and sintering rate would
lead to a lower degree of consolidation, higher porosity, and less
strength. On the other hand, semi-crystalline polymer powders
(e.g. PE, PP, PEEK, PA) have an ordered molecular structure with
sharp melting point. They do not gradually soften with a temper-
ature increase, instead, semi-crystalline materials remain solid un-
til a given quantity of heat is absorbed and then rapidly change
into a low viscosity liquid. The low viscosity of semi-crystalline
polymers above their melting temperature (T
m
) favors the rate and
amount of consolidation and gives rise to the parts with full den-
sity. Therefore, the density and mechanical properties of the fi-
nal part are comparable with the samples prepared by conven-
tional methods [62] . As such, for manufacturing a 3D structure,
semi-crystalline thermoplastics polymers are preferable to amor-
phous polymers [ 26 , 60 , 63 ]. However, crystallization rate should be
kept relatively slow (relative to the vertical build rate) in order to
avoid part distortion due to freezing shrinkage [64] . As such, the
low consolidation shrinkage of amorphous polymers in SLS result
in high dimensional accuracy of the final laser-sintered parts and
make them suitable candidates for producing mold for molding or
casting. Thus, it is important to evaluate the characteristics of the
powder used in laser sintering by different methods such as scan-
ning electron microscopy, density, thermal analysis, hot stage mi-
croscopy, and X-ray diffraction (XRD) analysis.
SLS can process polymers with only one fixed controlled con-
solidation temperature or several polymers with very close tem-
perature peaks with typically 5-10 °C difference. This indicates that
polymers with more than one tight melting range are not process-
able with SLS [64] . Besides polymer structure, the mean molecular
weight (MW) is another important property of a polymer which
ensures its processability with SLS. Some studies have reported a
better laser sinterability of poly- ε-caprolactone (PCL) with MW of
50.0 0 0 g/mol than PCL with a MW of 40.0 0 0 g/mol, due to big dis-
tortions caused by shrinkage [64–66] . This shrinkage is caused by
the lower molecular weight and the purity differences.
3.2. Metals
Metallic biomaterials are the most widely used materials for
orthopedic and dental applications. Their high yield strength,
high ultimate tensile strength, high fatigue resistance and fracture
toughness, along with fair ductility have all made metals the mate-
rial of choice for traditional load-bearing applications [ 13 , 67 ]. The
rise of metal AM in the field of biomaterials has been significant.
AM of metals has its roots in the stereo-lithographic process in-
vented by 3D systems [5] . The transition of AM from plastics to
the metals arena can be traced back to as early as 1994 when
EOS’s DMLS technology launched its metal powder-based proto-
type for a commercial system [ 68 , 69 ]. According to the ASTM clas-
sification, AM technologies for metals can be broadly classified into
two categories, DED and PBF. The PBF technologies enable building
of complex features, hollow cooling passages, and high-precision
parts that are the required features and specifications for building
load-bearing biomaterials. However, they are limited by the build
envelope and a single material per build. On the other hand, the
DED technologies offer a larger build envelop and higher deposi-
tion rate, while their ability to build hollow cooling passages and
finer geometry is limited [5] .
The requirements for metal powders for AM are more stringent
since the size, shape, and material chemistry are critical for a suc-
cessful and reproducible process [43] . The chemical composition of
metals can affect the melting temperature, mechanical properties,
weldability, thermal properties, among others. Due to the involve-
ment of fusion in L-PBF technology, the weldability and castability
of metal powders appear to play more important role in the qual-
ity and performance of the final part. For alloys such as copper,
aluminum, silver and gold that have a high reflectivity (hence low
absorption) and high thermal conductivity, the formation of an ef-
fective melt pool is difficult. Thus, higher power lasers along with
different wavelengths are recommended to increase the laser ab-
sorption. On the other hand, the size of powders can affect the
amount of consumed energy density to reach a certain density-. -
Both the size and morphology of powders can also significantly af-
fect the layer thickness and surface roughness. Fig. 3 shows spher-
ical powders of Ti, tantalum (Ta) and their mixture used for the
SLM processing of Ti–25Ta alloy [64] .
Although the metal powders used by PBF systems are similar
to those used in some conventional manufacturing processes such
as powder metallurgy and metal injection molding, there also exist
significant differences, especially when complete powder melting–
solidification occurs during the process (e.g., SLM and EBM). AM
metal powder has only captured a small fraction of the total metal
powder market. For most traditional designs, the metallic materials
fabricated via AM should possess mechanical properties compara-
ble to or slightly better than those of wrought parts, which is gen-
erally believed to be associated with their refined microstructure
[58] .
Along with stress-relieving via heat treatments, the level of
density has a significant effect on the mechanical properties of the
final parts [71] . For AM of metal parts, the SLM method typically
produces parts with more than 90% density under the optimized
process parameters. Nonetheless, the level of density is also de-
pendant on the type of metal in use. Ti and cobalt–chromium–
molybdenum (Co–Cr–Mo) alloys can be melted more easily via
SLM to almost full density than ferro powders [72] . In the DMLS
method metal powders are partially solid-state sintered together.
Therefore, to obtain densities greater than 90%, post-processing op-
erations such as sintering (e.g. HIP) and liquid metal infiltration
are used [73] . Load-bearing metallic biomaterials command a high
price and fit well with PBF technologies. Nonetheless, the range
of materials is still limited in PBF and included mainly stainless
steels, Co–Cr alloys, and Ti-based alloys [9] .
Steel is often used in many applications due to its strength and
thus researchers have carried out numerous studies regarding the
feasibility of using AM steel parts for load-bearing biomedical ap-
plications [74–78] . Most of these studies have been focused on
austenitic 316L and also martensitic grade AISI 420. It has been re-
ported that the mechanical properties of metallic parts of the same
alloy can vary significantly based on the AM process.
As opposed to 316L stainless steels, Co–Cr–Mo alloys exhibits
higher corrosion resistance and much lower allergic reaction with
human tissues, while it possesses higher wear resistance than Ti
alloys in tribological contact situations [79] . Ti is the most pre-
ferred load-bearing implant material worldwide due to its high
biocompatibility, high mechanical properties, and excellent corro-
sion resistance. Cp-Ti (ASTM 67) and some Ti alloys, -by virtue of
their excellent mechanical, physical, and biological performance,
are typically used in L–PBF to produce tailor-made load-bearing
implants [ 80 , 81 ]. For instance, the Ti–6Al–4V alloy alone is respon-
201

A. Nour i, A. Ro hani Shirvan, Y. Li et al. Journal of Materials Science & Technology 94 (2021) 196–215
Fig. 3. SEM micrographs of: (a) Ti powder, (b) Ta powder, and (c) Ti–25Ta powder mixture used for SLM process. The powder mixture shows a uniform distribution of the
Ta powders (brighter) within the Ti powder (darker) [70] . Reproduced with permission from Elsevier.
sible for consuming more than 50% of all Ti produced worldwide
and it is the main alloy used in L–PBF processes [52] . SLM method
has also been used for manufacturing some newly-developed Ti-
based alloys and composites, such as Ti-Nb alloys [ 82 , 83 ], Ti-Ta al-
loys [84] , and Ti/hydroxyapatite (HA) composites [85] .
In recent years L–PBF technology has also been used to pro-
duce load-bearing biomaterials from other metals such as nickel
titanium (NiTi) shape memory alloys [ 86 , 87 ], magnesium (Mg) al-
loys [ 24 , 88 , 89 ], zinc (Zn) alloys [90] , Ta [91] , and high-entropy al-
loys (e.g. TiNbTaZrHf) [23] .
4. Load-bearing applications of L–PBF biomaterials
The first reported use of additively manufactured biomateri-
als in orthopedics dates back to 1998, when these materials were
used for producing guides in spinal surgery to aid in the place-
ment of pedicle screws. Since then, 3D-printed biomaterials have
been used for shoulder, hip, knee, and ankle elective and trauma
surgery [92] . The most important load-bearing applications of L–
PBF manufactured biomaterials include arthroplasty, cranioplasty,
reconstruction of bone defects, dentistry, and orthopedic fixation
devices.
Bone-tissue engineering has gained a great attention in recent
decades due to its potential in repairing damaged or lost bone tis-
sues caused by severe accidents or weak bones (i.e., osteoporo-
sis). Osteoporosis is a common disease which occurs when the
bone becomes very weak and fragile and more likely to fracture,
in particular at the pelvis, hips, wrists and spine [93] . Therefore,
bone-tissue engineering addresses these problems to stimulate the
timely and effective regeneration of damaged bones. Bone scaffolds
provide an initial platform for cell attachment, proliferation and
tissue regeneration. Generally, bone strength is influenced by dif-
ferent parameters such as bone geometry (bone shape and size),
bone mineral density, microarchitecture, and degree of mineraliza-
tion.
Manufacturing of structural scaffolds and implants through L–
PBF for load-bearing applications such as orthopedic fixation de-
vices, cranioplasty, and hip/knee arthroplasty has opened new
prospects in clinical medicine. These AM methods provide the op-
portunity to design various shapes and sizes of scaffolds with dif-
ferent percentages of pores and mechanical strength to prevent the
stress shielding effect [94] . An ideal bone scaffold should be bio-
compatible, biodegradable, with a suitable mechanical properties,
and possesses high porosity and interconnectivity between pores
[ 95 , 96 ]. Furthermore, for adapting the bone’s geometry to the new
loading conditions, the structural scaffold should enhance the os-
teogenic healing process, especially in the case of large bone de-
fects [97] .
The most common technique in L–PBF for manufacturing
biopolymers is SLS. A combination of natural/synthetic biopoly-
mers with ceramics embraces the good ductility of biopolymers
and high bioactivity of ceramics. Therefore, a combination of nat-
ural and synthetic polymer-ceramic composites can be a promis-
ing option with a high strength and bioactivity, particularly for
structural applications. Moreover, a number of structural metallic
implants for the skull, hip, knee, elbow, shoulder, and jaw have
been successfully fabricated by SLM. As previously discussed, these
metals are stainless steel, Ti and its alloys (especially Ti–6Al–4V),
and Co–Cr alloys. As well as orthopedic applications, dental im-
plants fabricated by SLM are the other potential applications of
load-bearing biomaterials [94] .
Craniofacial problems are other important types of bone-related
defects that cause soft tissue or bone deficits. Due to the impor-
tance of esthetic outcomes and the number of tissues and struc-
tures in the craniofacial region, treatment of these types of de-
fects is very difficult. Grafting, local tissue rearrangement, micro-
surgical tissue transfer, and vascularized composite allotransplan-
tation are some traditional surgical treatments for craniofacial de-
fects. However, donor-site morbidity and procuring sufficient donor
tissue with the same properties are the most important problems
associated with these conventional methods [98] . Therefore, there
is a need for patient-specific treatments that provide structural
and functional replacement. To this end, AM technology can create
highly complex craniofacial geometries for specific needs [99] . In
craniofacial bone reconstruction, the scaffold must provide struc-
tural support for resident cells as well as transduction of mechan-
ical loads through the craniofacial skeleton [100] .
Computer-aided surgery has also been used to improve the ac-
curacy of implants in hip arthroplasty. Arthroplasty including hip
and knee joint replacement is a common surgical procedure that
has excellent long-term results [92] . The proposed potential bene-
fits of patient-specific instrumentation in total knee arthroplasty
include improving the accuracy of the bone cuts and thus the
alignment of the implants and improving the operative time and
efficiency [101] . Thus, the concept of patient-specific implants has
gained great attention in this area. These types of AM technolo-
gies use personalized algorithms to manufacture implants that ac-
curately resemble the patient’s original joint before the onset of
arthritis. Although the materials used in AM for arthroplasty im-
plants are limited, the use of AM methods is estimated to grow
significantly [102] .
The main advantages and disadvantages of L–PBF techniques for
the fabrication of structural implants are summarized in Table 2
[ 101 , 10 3 ].
4.1. Orthopedic and traumatological applications
The aging of the population and accidental bone injuries have
resulted in a continuing increase in various bone diseases. Ortho-
pedic surgery is used as the most fundamental treatment to repair
or fix bone fractures/defects. Orthopedic implants and prostheses
202

A. Nour i, A. Ro hani Shirvan, Y. Li et al. Journal of Materials Science & Technology 94 (2021) 196–215
Tabl e 2
Some of the benefits and drawbacks of L–PBF technologies for the fabrication of load-bearing biomaterials.
Advantages Disadvantages
•Tailoring the microstructure and properties based on
individual patient needs, such as adjusting the
density of the implant to match bone density
•Production of bone implants with biomimicry
features
•A definable degree of surface roughness that helps
bones and implants to fuse better
•Lattice structures that can help to accelerate
post-operative healing significantly
•Shortening of hospitalization period and minimizing
unpleasant side effects
•Reducing aseptic loosening to delay revision surgery
•Cost is a major factor as specialized machinery costs
a lot. While AM by itself consumes less time,
prefabrication and post-processing may be intensive
and are not straightforward. Furthermore, the
associated postproduction equipment takes up at
least as much space
•The fine metal powders and even finer nanoparticle
waste represent a significant health and safety
challenge
•The high thermal gradients present during SLM are
due to the short interaction times; the extremely
high heat input gives rise to residual stresses,
elemental segregation, stress cracking, formation of
nonequilibrium phases, and entrapped spherical gas
bubbles
•A large degree of shrinkage and warpage tends to
occur during liquid– solid transformation, decreasing
the dimensional accuracy and accumulating
considerable stresses in SLM-processed components
•PBF leads to lower surface quality than other AM
techniques due to the presence of large and partially
melted powder particles in the printed pieces’
surfaces
•Limited component size: the size of components is
restricted by the size of the build chamber
are applied to procedures from replacement of injured and dis-
eased hard tissues to trauma surgery [104] . The majority of im-
plants are available in average standard sizes that are designed
to fit for most patients. One can colloquially name these tradi-
tionally manufactured implants as the one-size-fit-all . However, pa-
tients with extreme anatomical variations and deformities may re-
quire specialized implants to ensure a proper fit. This is where
AM emerges to play a key role in orthopedics, traumatology, and
dentistry. Using this new technology, computer-aided design (CAD)
models of patient anatomy generated through digital imaging tech-
niques such as microcomputed tomography ( μCT) and magnetic
resonance imaging (MRI) allow for design and fabrication of cus-
tomized implants, also known as patient-specific implants. 3D-
printing models can help to assess the surgical approach for cor-
rective osteotomies, as well as the visualization of traumatic situa-
tions, such as complex bone fragmentation.
Orthopedics, traumatology, and maxillofacial surgery have been
among the first medical fields to use AM technology to build cus-
tomized or personalized implants. The use of customized implants
for surgical reconstruction has been shown to reduce the duration
of surgery. For surgical reconstruction of a large cranial defect, Jar-
dini et al. [105] fabricated a customized SLM Ti–6Al–4V implant via
obtaining a series of μCT data for extraction of the cranial geom-
etry. While it took only 3 h to place the customized implant into
the patient, the duration of surgery for non-customized implants
in similar surgery was approximately 6 h. The implant customiza-
tion eases the fabrication of orthopedic implants whose curves and
structural complexities cannot be readily produced by conventional
methods. Fig. 4 shows a customized Ti–6Al–4V bone plate fabri-
cated by SLM through customized design to satisfy the complex
requirements of different patients suffering pelvic fracture.
The advent of sterilizable and autoclavable materials for 3D
printers has also paved the way to the fabrication of personal-
ized implants for orthopedic and traumatology surgery [3] . Apart
from tailoring the prosthesis based on a specific anatomy, ortho-
pedic implants also need to integrate (or even regenerate) with
the patient’s own bone, which creates tissue support and prevents
implant failure [ 3 , 7 ]. AM technology is also capable of recreat-
ing an exact surface finish or a controlled porosity able to aid os-
teointegration. Highly porous implants can be created using high-
resolution AM techniques so as to secure the implant to the sur-
rounding bony tissue. This aspect is further discussed in the fol-
lowing sections.
Similar to other biomaterials for hard-tissue engineering, load-
bearing orthopedic biomaterials processed by AM should meet re-
quirements including biofunctionality, suitable mechanical proper-
ties, appropriate morphological structures, and structural process-
ability [17] . Implant materials should exhibit adequate biocompat-
ibility to ensure no inflammation or rejection by the surround-
ing environment. Surface modification and coating are effective
strategies to provide additively manufactured implants with desir-
able osseointegration properties and further improve their biolog-
ical functionality [107] . The mechanical properties of AM implants
such as their elastic modulus and strength should also be compat-
ible with those of bone. Unlike polymeric materials, the mechani-
cal properties of most metals used in AM are higher than the me-
chanical properties of bone. This leads to a mismatch in modulus
of elasticity and causes the stress shielding, which can be easily re-
solved by using porous structures and changing of the composition
of implants [ 14 , 108 , 109 ]. For a stable support function, it is also a
design requisite to ensure that the shape and outer contour of the
implant match the defect portion of the bone. Moreover, due to the
processing parameters and technical limitations of AM equipment,
the structure of customized metallic implants should be designed
and modified in such a way that assures the machinability of the
implant [17] .
Joint arthroplasty is the replacement of one or both sides of a
joint with load-bearing components. The most common reasons
are degenerated cartilage and subchondral bone associated with
osteoarthritis. Hip and knee arthroplasties are the most frequent
joint surgeries, followed by the shoulder, ankle, and other joints.
Although conventional prostheses have been commonly applied in
joint arthroplasty over the last decades, these prostheses are not
be a perfect choice for surgeons for severe bone defects around
joints. There are subtle differences in the anatomy of joints be-
tween each individual which prevent the conventional implants
203

A. Nour i, A. Ro hani Shirvan, Y. Li et al. Journal of Materials Science & Technology 94 (2021) 196–215
Fig. 4. The importance of customization for manufacturing a pelvic bone plate: ( a ) a customized Ti–6Al–4V bone plate after heat treatment and anodization; and ( b )
anatomical match between bone plate and pelvic model [106] . Reproduced with permission from MDPI.
Fig. 5. (a) External and (b) internal parts of an SLM-fabricated Ti–6Al–4V (ELI) acetabular cup, post-processed using HIP.
from meeting the requirements of patients. As such, a customized
AM prosthesis is an optimal candidate in this regard [49] . For in-
stance, based on the shape of the bone defect in a given patient,
more accurate and complex shapes of acetabular cups can be man-
ufactured using L–PBF in order to create a better match between
the cage and the host bone [49] . Fig. 5 shows the external and in-
ternal parts of a Ti–6Al–4V (ELI) acetabular cup fabricated by SLM.
The customization of the acetabular hip cup included the unique
placement of holes to permits patient-specific screw trajectory. On
this occasion, the recipient was a cancer patient suffering bone
loss.
Alongside metals, biopolymers are also playing an important
role in the manufacture of orthopedic implants via L–PBF. In com-
parison to conventional implants such as Ti, PEEK has an elastic
modulus comparable to that of cortical bone which significantly
reduces the stress shielding effect. Moreover, PEEK is radiolucent
which provides an MRI compatibility. It can also be sterilized by
radiation or steam-autoclaving without any changes in its prop-
erties. All these characteristics make PEEK a promising choice for
manufacture of orthopedic applications such as medical fixation
devices and implants used in cranioplasty, spine cages, and den-
tistry [ 110 , 111 ]. Using SLS, Tan et al. [112] developed a PEEK/HA
biocomposite scaffold for the regeneration of bone defects. First,
different mixtures of PEEK/HA powders were prepared by physi-
cally blending pure PEEK and HA via a roller-mixer. Addition of
HA, as a bioactive ceramic, can increase the bioactivity of the fi-
nal fabricated implants. Since PEEK has a lower melting point than
HA, it is possible to sinter it at a temperature near its T
g
at around
143 °C to bind the HA particles within the sintered PEEK matrix.
The study showed a promising result on laser sintering of PEEK, as
a high melting point polymer, in a lower temperature environment
as well as the usefulness of PEEK/HA composite for the regenera-
tion of bone defects.
Rimell et al. [113] studied the SLS process of UHMWPE compo-
nents for orthopedic applications. They characterized the capabil-
ity of SLS in the formation of continuous solid bodies and prob-
able chemical or physical changes in UHMWPE during the lasing
process. The obtained data indicated that SLS has the potential to
form continuous dense bodies, but only with the use of commer-
cial systems and specialized powders. The lasing process degraded
the UHMWPE through different mechanisms including chain scis-
sion, crosslinking and oxidation. However, it did not degrade the
properties to an unacceptable level for some applications. It is,
however, important to determine the extent of degradation when
the SLS system is used for manufacturing of small joints or den-
tal implants. In a patent, particles of UHMWPE and/or high-density
polyethylene (HDPE) and/or PP were fused together layer-by-layer
by means of a SLS. The produced implants showed the desired in-
growth of soft and bone tissues, as well as good durability and load
capacity. By using this technique, different load-bearing implants
with the most complex geometries for the skull, hand, sternum,
foot, and other structural areas can be produced [114] . In another
patent, a selected region laser sintering and molding method for
manufacturing of UHMWPE was introduced in three main steps:
1) heating the polymer powder in a selected laser-sintering re-
gion and molding to a pre-heated temperature; (2) molding the
UHMWPE powder by adopting a pre-set laser scanning speed and
output power; and (3) taking out the molded piece and keeping
the heat at a temperature of 110-130 °C for 8-12 h and then cool-
ing to obtain the molded part of polymer. The manufactured com-
ponent exhibited a high mechanical properties, high dimensional
precision, and small shrinkage, suitable for manufacturing medical
auxiliary tools [115] . Changhui et al. [116] used SLS to produce a
customized UHMWPE tibial insert for femoral component and tib-
ial trays and evaluated its mechanical properties and dimensional
accuracy. The obtained results showed that the tensile strength and
204

A. Nour i, A. Ro hani Shirvan, Y. Li et al. Journal of Materials Science & Technology 94 (2021) 196–215
Fig. 6. Radiographic images of a cemented total knee arthroplasty in the right knee
of a 73-year-old woman, six months after the implantation of the prosthesis. The
arrows indicate the PMMA bone cement [118] . Reproduced with permission from
MDPI.
elongation of the tibial insert increased from 14.1 to 24.1 MPa and
from 5.4% to 390%, respectively, by a post-heat treatment.
PMMA is also a suitable material for fixation of prosthetic im-
plants and repair of vertebral fractures and is frequently used as
implants for orthopedic applications and bone cements. However,
poor osteointegration and low bone attachment are the most im-
portant disadvantages of PMMA implants. Bioactive materials such
as bioceramics are usually used to overcome these problems and
to increase the mechanical and biological properties of this poly-
mer [117] . Fig. 6 shows the X-ray images of a PMMA bone cement
used in the total knee arthroplasty in the right knee of a 73-year-
old woman at six months follow up [118] .
Velu et al. [119] used SLS for the fabrication of a PMMA-based
composite for bone repair and reconstruction. For these applica-
tions, several forms of calcium phosphates as a bioactive ceramic
in combination with biopolymers are often used. PMMA and β-
tricalcium phosphate ( β-TCP) were introduced as a suitable bio-
composite for specific medical applications. The results indicated
that laser power and scan speed have great effects on the interpar-
ticle fusion, coalescence, and sintering. It was also shown that the
inter-particle coalescence varied significantly with varying process
parameters including laser power and scan speed. Increasing the
concentration of β-TCP from 10% to 20% changed the morpholog-
ical properties of the biocomposite. The thermal aspects of SLS of
PMMA/ β-TC P composites were also studied by the same research
group. Photomicrographs of samples sintered with varying process
conditions showed different morphologies. These variations were
attributed to changes in the thermal fields because of using differ-
ent amounts of the bioceramic filler. Moreover, the results showed
that SLS is an effective tool in building well-coalesced layers at a
low concentration of β-TCP ( < 10 %) and lower laser power (~38 W).
However, higher β-TCP concentrations led to greater porosity and
larger pore sizes [120] .
An interbody fusion cage (colloquially known as a "spine cage")
is an implant used in spinal fusion procedures when the disc space
between two vertebrae is narrowed or distracted. The objective is
to restore lost disc height resulting from a collapsed disc and to re-
lieve the pressure on nerve roots. These implants are either cylin-
drical or square-shaped, and provide a comfortable environment to
bone grafts for bony healing. The excessive cage stiffness of a con-
ventionally made cylindrical implant may result in several compli-
cations, including the movement of the cage, stress shielding, and
pseudoarthrodesis [ 121 , 122 ]. In spinal surgery, AM techniques play
a vital role in planning and designing of various spine cages for fu-
sion surgery or vertebral body replacement and disc implants for
total disc replacement [ 57 , 123 ]. Lin et al. [124] fabricated a Ti–6Al–
4V lumbar interbody fusion cage using SLM that could reproduce
Fig. 7. Medtronic’s ARTiC-L Spinal System. Courtesy of Medtronic PLC.
an intricate microscopic structure with a compressive modulus be-
tween those of cancellous and cortical bone (2.97 ±0.90 GPa). In or-
der to add extra support and strength to the fusion while healing,
pedicle screws are sometimes used in spinal fusion. Burnard et al.
[125] designed an additively manufactured implant with fixation
holes for pedicle screws which was successfully applied in recon-
struction of a Ti primary bone tumor. In another study, a porous
Ti–6Al–4V cage with macro- and micro-architecture was produced
by SLM [126] . The effects of additive angle on surface properties
and biocompatibility of dense Ti–6Al–4V discs with six different
angles were investigated. With increase in additive angle, the num-
ber of unmelted metallic particles increased and thus resulted in
increase in both surface roughness and hydrophobicity. This study
indicated the great potential of SLM-fabricated porous Ti–6Al–4V
cages for spinal fusion. Medtronic PLC has recently launched an
SLM Ti–6Al–4V spine cage, via TiONIC Technology, which enabled
more complex designs and enhanced surface textures for spine
surgery implants ( Fig. 7 ). As opposed to smooth materials, a rough
surface texture has been shown to promote osteoconductivity and
bone response.
PEEK cages have been widely used during the past decade
due to their biocompatibility, radiolucency, durability, strength, and
elastic modulus comparable to that of cortical bone [ 127 , 128 ]. It is
considered as a better candidate for spinal fusion than Ti. Never-
theless, the combination of PEEK with Ti and osteoconductive ma-
terials such as HA can improve osseointegration.
L–PBF-manufactured implants have also gained much interest
in orthopedic trauma. The word “traumatology” is a broad term to
describe all kinds of injuries that affect the bones, joints, muscles,
tendons, and ligaments in any part of the body that are caused
by accidents, sport injuries or violence to a person. This can range
from minor fractures to severely broken bones with a direct threat
to the patient’s life. Traumatology is considered a sub-specialty of
orthopedic surgery in which fractured bones are treated and the
injured parts of the body are restored to their original strength and
maximum function prior to the injury. Complex bone fractures and
large defect sizes are always tough problems. It was found that AM
implants were ideal candidates to heal defect sizes exceeding 8 cm
[49] .
The objective of fracture-fixation implants is to reduce the
bone fragments and restore their original anatomy. Plates and in-
tramedullary nails are the two typical types of fracture-fixation
devices. The screws are often similar in material to the main
plate to reduce galvanic corrosion [129] . Plates are positioned out-
side the bone cortex, whereas nails are inserted within the cortex
down the intramedullary canal of long bones. In these areas, AM
technology has mainly focused on patient-specific implant shape,
patient-specific screw position/orientation and the curvature of in-
tramedullary nails. Additively manufactured plates, and in partic-
ular patient-specific plates, provide different approaches for bone-
fragment stabilization, match better to bone shape, and assist with
obtaining proper reduction without requiring intraoperative plate
contouring [ 130 , 131 ]. It is, however, worth noting that for plate
205

A. Nour i, A. Ro hani Shirvan, Y. Li et al. Journal of Materials Science & Technology 94 (2021) 196–215
Fig. 8. Application of SLM process in craniomaxillofacial surgery. Grade 23 Ti–6Al–4V (ELI) cranial and mandibular implants for: (a) mandibular resection; (b) fixation of
mandibular angle; (c) fixation of cranial bone flaps with a patient-specific cranial implant; and (d) custom mandible surgical guides. Figs. 8 (a-c) courtesy of Bonash Medical
and Fig. 8 (d) courtesy of Adeiss centre, Wes tern University.
fixation it is important to insert screws into denser, higher qual-
ity bone to reduce the risk of screw loosening, especially for os-
teogenic or osteoporotic patients. In patients with preexisting joint
replacements near a fracture, it is necessary for screws to avoid
the prosthesis [132] . In more complex fractures, it is possible that
a screw will unintentionally passes through the fracture site, and
hence, reduces mechanical stability and impairs healing. AM cus-
tomized intramedullary nails for long bones such as the femur
can better match the anatomical curvature of the patient’s in-
tramedullary canal [133] . Non-anatomically conforming nail radii
can result in painful pressing of the nail on the inner cortex, injury
during nail insertion, or angular deformity in the limb as the bone
conforms to the nail through motion at the fracture site [129] .
4.2. Craniomaxillofacial applications
AM technology is also extensively used in maxillofacial
[134] and craniofacial [135] surgeries. Although high mechanical
loads are not generally applied on cranial bone, the mechanical
properties of a material used in cranioplasty have to associated
with the anatomical location due to the variance of bone thickness
and thus need to be superior to those of skull bones. The need for
craniofacial/maxillofacial surgery may arise from different reasons.
In some cases, patients suffer fractures of the skull, chin and jaw
due to trauma. In other patients, cancer may spread to a region of
the jaw and the infected tissue needs to be removed. AM may be
applied to replace all or a portion of the infected tissue [136] . The
SLM and DMLS processes have enhanced the scope for fabricating
craniofacial and maxillofacial implants due to high strength and
hardness. Further, image-processing software favors the accuracy
of fit of the prosthesis for use in human mandibles [137] . Grade 23
Ti–6Al–4V (ELI) cranial and mandibular implants fabricated by the
SLM process are shown in Fig. 8 .
Berretta et al. [138] studied the effect of four manufacturing
build orientations including horizontal, oblique, vertical and in-
verted horizontal on the properties of L–PBF PEEK cranial implants.
The implants were characterized in terms of dimensional accu-
racy, weight, and mechanical properties. The results indicated that
the implants manufactured with horizontal and inverted horizontal
orientations had the highest compressive strength resistance and
showed the least deviation from the design model. In contrast, the
vertically manufactured implant had the lowest accuracy in design
and 70% lower first failure compared to the inverted horizontal im-
plant. Pimentel et al. [139] studied the variation of SLS technique
parameters to fabricate PEEK scaffolds and characterized their mor-
phology and roughness, and the thermal processing conditions. It
was shown that the fabricated scaffolds were suitable for cranio-
plasty applications due to their high surface roughness. Based on
the results, parameters such as powder bed temperature and laser
power should be controlled to reduce delamination at the cross-
sections of scaffolds. Additionally, fast cooling in the SLS process
can decrease crystallinity.
There are some strategies to increase implant fixation to sur-
rounding bone and to improve the bone-implant interface. For ex-
ample, topographical modification of an implant surface or the
combination of a bioinert implant with bioactive materials can en-
hance the bone-implant fixation. With increasing the osteointegra-
tion and biocompatibility, the bone-implant fixation can increase
and provide a better strength for load-bearing applications. In a
study, customized PEEK scaffolds were designed to incorporate a
trabecular structure using a CAD program and printed via the SLS
method; the combination of 3D-printed PEEK scaffolds with mes-
enchymal stem cells (MSCs) could overcome the limitations of us-
ing this polymer for structural bone regeneration in craniofacial re-
construction [140] .
Alongside PEEK, PMMA is another biopolymer that is fre-
quently utilized for cranioplasty with long-term results. Fiaschi
et al. [141] conducted a study of patients with cranial problems.
A patient specific PMMA implant was manufactured using the SLS
method. It was revealed that the manufactured implants had low
206

A. Nour i, A. Ro hani Shirvan, Y. Li et al. Journal of Materials Science & Technology 94 (2021) 196–215
Fig. 9. A biodegradable patient-specific cranial implant demonstrator made by SLS from PLA/CC composite powder on the customized Formiga P 11 0 machine [143] . Courtesy
of KLS Martin Group.
complication rates and the custom-made technique was associated
with an excellent level of patient satisfaction in the long term.
One of the most serious problems in processing of PLA with SLS
is the formation of a high level of micro-pores due to incomplete
coalescence of powder particles as a result of high melt viscosity.
This leads to poor mechanical properties and limits the application
of laser-sintered PLA. Therefore, inorganic biomaterials such as cal-
cium carbonate (CC) can be used to improve the cell reactions and
promote bone regeneration [142] . In a study, four PLA/CC compos-
ite powders with different inherent viscosity levels were made. The
synthesized composite materials with the lowest inherent viscosity
had the best processability by SLS. The small diameter of polymer
particles (50 μm) and small zero-shear melt viscosity (400 Pa.s)
gave rise to fast sintering. The SLS process parameters were de-
veloped to achieve low micro-porosity, good biocompatibility and
low polymer degradation. Thus, the prepared scaffolds have great
potential to be used as patient-specific bone replacement implants
[143] . Fig. 9 shows biodegradable patient-specific cranial implant
samples made by SLS of PLA/CC composite powder. The intercon-
nected pore structure was designed using Autodesk’s Within soft-
ware. The strut diameter was about 1 mm and the build process
took only about 2 h. These types of biodegradable patient specific
implants made by SLS method offer a perfect fit to the cranial de-
fect and can significantly improve the bone defects-related medical
treatments [144] .
4.3. Dental applications
Besides orthopedic and craniomaxillofacial applications, den-
tistry is another avenue for structural biomaterials that has greatly
benefited from AM technology. Creating artificial teeth and dental
implants has become progressively widespread and simplified with
AM technologies due to the ability of the process to create them
rapidly (time-to-market), sometimes within the dentist’s office it-
self [ 145 , 146 ]. Although there are several possible applications of
AM technologies in dentistry (e.g. dental restorations, models, sur-
gical guides, and orthodontic materials) [ 46 , 147 ], only the load-
bearing dental restorations including crowns, bridges, dentures,
and implants are the focus of the current review.
Traditionally, major parts of dental crowns and bridges are
manufactured manually based on lost wax casting, which is not
only a time-consuming process but also strictly depends on dental
technician skills [9] . Similarly, fabricating dental implants that re-
quire subtractive manufacturing methodologies (i.e., machining or
milling) may cause high material wastage and energy consump-
tion, hinder the ability to fabricate complex shapes, and require
expensive tooling and setup, although, milling/machining may be
used to refine the final printed form. Compared to conventional
methods and subtractive manufacturing, there are significant ben-
efits in fabrication of load-bearing dental restorations through AM.
AM technology: (i) produces dimensionally accurate parts; (ii) is
cost-effective; (iii) is a reliable process to control the depth and
width of teeth; (iv) is a unique method to fabricate customized
dental restorations and implants; (v) is rapid and time-saving; (vi)
is a digital storage-based method that considerably reduces the in-
ventory of physical models; (vii) limits the human error relevance
in the procedures by virtue of the lower number of manufactur-
ing steps; and (viii) reduces the environmental impact, ensuring
greater manufacturing sustainability [ 11 , 148 ].
AM of clinical devices using L–PBF processes has been em-
ployed successfully for the fabrication of dental crowns, dental
bridges, and dentures due to its ability to produce complex-shaped
objects rapidly in hard-wearing and corrosion resistant metals and
alloys directly from their CAD data, leaving only the final finishes
of restorations to be applied by hand [ 1 , 14 9 ]. Dental crowns may
be manufactured through SLS or SLM depending on the type of
crown [136] . In a study by Strub et al. [150] , the use of SLS to fab-
ricate both ceramic and metal dental restorations was described.
Averyanova et al. [9] produced dental crowns and bridges with ad-
equate density, geometrical accuracy, mechanical properties, and
microstructure. The SLM Co–Cr dental alloys showed higher me-
chanical, corrosion resistance, and tribological properties, compar-
atively good fitting ability, and higher adhesion strength in the
porcelain compared with the cast alloys [48] . Wu et al. [151 ] also
evaluated the mechanical properties, porcelain bond strength, and
metal-ceramic interface of an SLM Co–Cr dental alloy. It was con-
cluded that the SLM alloy exhibited improved mechanical proper-
ties and similar porcelain bond strength to the currently used con-
ventional cast alloy for metal-ceramic restorations. Similar to den-
tal crowns, metallic dentures have also been fabricated using L–
PBF. There are some studies in the literature that report the use of
Co–Cr–Mo alloys for the fabrication of metallic dentures [152–154] .
One of the other advantages of L–PBF in dentistry is that it allows
the fabrication of novel dental implants with a porous or rough
surface [ 155 , 156 ]. Tolochko et al. [157] utilized a combined SLS and
SLM process to fabricate dental implants with a graded structure
composed of a remelted compact core and a sintered porous shell.
Fig. 10 depicts a Ti–6Al–4V dental implant produced by DMLS pro-
cess. Acid etching was performed to roughen the surface and thus
improve the biological responses.
Like AM of orthopedic implants, special attention is also paid
to customized near-net-shape dental restorations such as crowns,
bridge frameworks, and partial denture frameworks [159] . As such,
these dental pieces can be manufactured with intricate details (e.g.,
irregular grooves, crannies, valleys) with no extra production cost,
since the product complexity is not be costly beyond the design
207

A. Nour i, A. Ro hani Shirvan, Y. Li et al. Journal of Materials Science & Technology 94 (2021) 196–215
Fig. 10. SEM micrographs of dental implants fabricated by DMLS: (a) as-fabricated
implant; (b) fabricated implant after etching process [158] . Reproduced with per-
mission from Elsevier.
stage. Once the design is set in AM technology, costs are indepen-
dent of the shape (i.e. a crown and a cube are processed in the
same way) [11] . Nonetheless, the consumed material in a given de-
sign also dictates the overall cost of the manufactured parts.
From the materials standpoint, the whole range of dental mate-
rials can be used in L–PBF for manufacturing of dental restorations
including metals and alloys (Ti and Ti alloys, Co-Cr alloys, stain-
less steels), thermoplastic polymers and composites, waxes, and
ceramics [ 39 , 48 ]. Among this wide range of materials, AM of ce-
ramic dental materials is still underdeveloped, mainly due to the
difficulty of manufacturing components with suitable surface fin-
ishes, mechanical properties, and dimensional accuracy [11] . It is
notable that dimensional accuracy is critical between the digital
model and the printed construct, because the manufactured parts
must fit tightly to meet the needs of each patient. Assessment of
the accuracy and reproducibility of dental models usually consists
of measurements of intercanine distances, intermolar distances, the
overjet, the overbite, tooth sizes and arch lengths [46] .
There are a number of factors that affect the dimensional accu-
racy. The work of Lee at al. [160] summarize the dimensional accu-
racy in different manufacturing processes. The work describes the
lower accuracy and lesser improvement in the Z -direction than in
the other directions ( X and Y ), since the former direction is influ-
enced by a variety of process parameters that cannot be controlled
easily, e.g., nonuniform compaction/densification of the powder
within the layers, evaporation of material by laser and shrinkage
during solidification. In L–PBF processed components, the possible
formation of low melting porosity between the layers decreases
the effective load-bearing area perpendicular to the layers (in the
Z -direction) and gives rise to stress concentration (the notch ef-
fect). This leads to reductions in static and dynamic strength in
the Z -direction [161] . If feasible, minimizing the Z -direction/height
would not only resolve the above issues to a great extent but also
expedite the build time and minimize the powder usage.
With the existing drawbacks of L–PBF technology in dentistry,
it is not clear whether 3D-printed dental components will cap-
ture the market in the near future and overtake their conventional
counterparts. While we still do not know which AM technology
will dominate the market, it is fair to state that these visions are
not absolute and are defined on a case-by-case basis. L–PBF tech-
nology in dentistry is in its infancy and thus, long-term clinical
studies are necessary to examine the accuracy, reproducibility, and
safety of 3D-printed dental components [46] .
5. Porous scaffolds in load-bearing applications
Bone is a complex porous composite structure with specific
morphology and mechanical properties such as viscoelasticity and
anisotropy [162] . Cancellous bone (also known as trabecular or
spongy bone) is located in the core of bones and contains a large
number of interconnected pores (i.e., 50–90%) and fairly equiaxed
pore sizes with a diameter of hundreds of micrometers. The high
porosity of cancellous bone results in significantly lower mechan-
ical properties compared to those of cortical bone. As discussed
earlier in this review, a mismatch of elastic modulus between load-
bearing implants and natural bone may give rise to stress shield-
ing, and consequent loosening and failure of the implant. There-
fore, a porous scaffold material with low elastic modulus that does
not cause the stress shielding effect is a requisite for successful im-
plantation [14] .
Bone replacement constructs including porous scaffolds for
hard-tissue engineering need to be biocompatible with surround-
ing tissue, radiolucent (a minimal MRI artefact), easily shaped or
molded to fit perfectly into the bone defect, nonallergic and non-
carcinogenic, exhibit sufficient strength and stability over time, and
be able to support bone growth (osteoconductivity) and preferably
encourage/stimulate the ingrowth of surrounding bone (osteoin-
ductivity) [162] . In almost all clinical cases, the scaffolds f or phys-
iological load-bearing areas are designed to be permanent. They
must retain their shape, strength, and biological integrity through
the modeling and remodeling processes of the damaged bone tis-
sue.
The implantation of porous scaffolds has been a major advance-
ment in the field of orthopedics within the past few decades. Nu-
merous studies have been carried out on the design of porous scaf-
folds [163–166] . The ideal bone scaffold is found to have three
characteristic features: proper surface roughness; suitable porosity
(permeability); and appropriate mechanical properties. It has been
demonstrated that rough surfaces are more osteoconductive than
smooth surfaces and hence can promote better and faster bone
apposition. SLM components typically have micrometer roughness
in the range of 2–20 μm [50] . This is attributed to the thermal
gradients across layers and the staircase effect [167] . Wong et al.
[168] and Carlsson et al. [169] reported that a surface with rough-
ness between 1 μm and 10 μm is optimal for promoting bone
apposition, stimulating mesenchymal cell differentiation into func-
tional osteoblasts. In order to have a margin of security against
risks, it is imperative for a manufactured scaffold to exhibit an
equal or excess strength to ensure an equivalent or better load-
bearing capability than natural bone. Last, it is very important to
consider the fatigue strength of a non-biodegradable scaffold, since
a long-term scaffold will be exposed to cyclic loadings during the
patient’s life.
Although high porosity may offer space for bone ingrowth
without the aid of additional coating and so benefit implant fix-
ation, the strength and ductility of porous structures are reduced
with increasing porosity. The mechanical properties of the porous
scaffolds are very sensitive to final density, pore size, material
type, and fabrication parameters [ 170 , 171 ]. Hence, it is necessary
to maintain a trade-off between porosity and mechanical proper-
208

A. Nour i, A. Ro hani Shirvan, Y. Li et al. Journal of Materials Science & Technology 94 (2021) 196–215
ties, since the relationship between biomaterial microarchitecture
and mechanical properties has a potential impact on the biologi-
cal performance of load-bearing implants [ 49 , 17 2 ]. A material with
gradient porosity from high to low could lead to a gradual mechan-
ical property variation from low to high. Therefore, the fabrication
of functionally graded materials (FGMs) could be an ideal option
to maintain a trade-off between porosity and mechanical strength
[ 165 , 173 , 174 ]. Yan g et al. [175] investigated the use of triply pe-
riodic minimal surface (TPMS)-based structures with the continu-
ity along the outer surface to improve the additive manufactura-
bility and diminish local stress concentrations of strut-based cel-
lular structures. The SLM manufactured TPMS structures exhibited
high permeability, specific surface area (surface to volume ratio),
and specific strength (strength to weight ratio). In another study
by Hang et al. [176] continuous functionally graded porous scaf-
folds (FGPSs) based on the Schwartz diamond unit cell were man-
ufactured by SLM with a wide range of graded volume fraction
from 7.97 % to 19.99%. The scaffolds exhibited no defect with a good
geometric reproduction of the original designs. By adjusting the
graded volume fraction, the Young’s modulus and yield strength
of the Ti diamond FGPSs can be tailored to those of the cancellous
bone.
There are technical limitations in conventional methods which
undermine their ability to precisely control pore size, pore geom-
etry, pore interconnectivity, spatial distribution of pores, level of
porosity, etc. As a result, there are very few manufacturing tech-
nologies capable of producing porous structures that possess the
majority of the desired requirements [162] . Similarly, the manu-
facturing of FGMs with precise gradient and designed porosity is
difficult using conventional methods such as casting, powder met-
allurgy, and injection molding. The advent of AM has opened up
a host of opportunities to fabricate complex implants with a con-
trollable pore structure and optimal properties as sophisticated as
those of cancellous bone. The porous metallic and polymeric scaf-
folds manufactured by L–PBF process could be an optimal choice
in load-bearing applications which can solve the drawbacks of tra-
ditional scaffolds. Given these technical capacities, AM is providing
the ability for the design of FGM implants such as fracture-fixation
devices [177] . Full, partial, or gradient porous structures can not
only minimize stress shielding, but also promote bone in-growth
for implant fixation. In addition, AM porous scaffolds can shorten
surgical time due to the customized design [ 77 , 17 8 ]. The revolu-
tionary precision of these techniques makes it possible to join very
thin sections (from 20 to 60 μm) together and permits very com-
plex porous geometries with thinner struts [158] . Like traditional
scaffolds, L–PBF scaffolds are also fabricated using different mate-
rials including metals, polymers, ceramics or their composites in
which more than one class of material is employed in the scaffold.
5.1. Porous metallic scaffolds
One of the most common methods to create porous and rough
textures on load-bearing biomaterials is coating. Porous metal-
lic coatings feature high surface areas beneficial for osseointe-
gration and biological fixation, as well as a microarchitectural
pore interconnectivity facilitating body fluid and nutrient transport
[ 179 , 18 0 ]. However, porous metallic scaffolds go beyond coatings
when load-bearing capacity is required in which the porous fea-
tures that promote osseointegration exceed the surface and extend
throughout the body of the implant [53] .
Technologies to fabricate porous biomaterials with a controlled
microarchitecture are crucial to the success of load-bearing scaf-
folds. Among other things, the extreme chemical affinity of many
biocompatible metals (e.g. Ti and its alloys) to atmospheric gases
such as oxygen, hydrogen, and nitrogen, significantly deteriorate
the ductility of porous metallic scaffolds [181] . In view of ma-
terial properties, some metals are either too weak to be shaped
into the desired architecture with a controlled porous structure or
too stiff and would fracture upon arrangement into certain archi-
tectures [162] . From a manufacturing point of view, conventional
manufacturing methods are neither easy to employ nor successful
in producing such complex porous structures within the tight con-
straints of porosity, optimum pore size, and mechanical strength
that are required [182] . To this end, AM technology, and in partic-
ular L–PBF, is a promising manufacturing technique to overcome
the existing drawbacks of conventional methods in fabrication of
porous metallic scaffolds for load-bearing applications [183] . The
application of L–PBF processes for fabricating 3D porous metallic
structures has pushed the development of designing pore config-
urations, interconnectivity, size and distributions to a new phase
and made a breakthrough progress. These 3D architectural struc-
tures are composed of consecutively and repeatedly arranged or
random pores that are preferably designed to be interconnected for
load-bearing biomaterials. There are distinctive definitions with re-
spect to orientation and repeatability of pores in the given space.
For instance, honeycomb structures are two-dimensional cellular
structures identified with unit cells with the same size and shapes
including tetrahedron, triangular prism, square prism, and hexago-
nal. On the other hand, the shape of unit cells in the foam struc-
tures is randomly generated and the cell walls have random orien-
tation [ 184 , 185 ]. Subsequently, lattice structures are defined as the
3D structures composed of consecutively and repeatedly arranged
interconnected cells [186] .
There is also another classification for the unit cell topol-
ogy, known as parametric or nonparametric design. The cellu-
lar structures in parametric design are generated according to
specific algorithms, whereas nonparametric design is not consti-
tuted by algorithm. There are two main methods to design the
porous structures according to algorithms (i.e. parametric), namely
Voronoi-Tessellation, and TPMS. Nonparametric design, also called
geometry-based design, is referring to structural and geometric de-
sign that comprises the diamond, the body-centered cubic (BCC),
and the polyhedral structure. The topology of unit cell is one of
the key factors that determines the mechanical property of porous
scaffolds. Chen et al. [187] carried out an excellent review on
porous scaffold design for AM in hard tissue engineering applica-
tions. Han et al. [188] designed the BCC unit cell to construct the
3D printed porous Co-Cr-Mo bone substitutes. The construct was
ascribed to octahedron, containing a node located at the center of
a cube, where all the struts radiate out to the corners of the cube.
In another study, porous Co-Cr-Mo alloy scaffolds were fabricated
by SLM in four different unit cell topologies, i.e., BCC, cubic close
packed (CCP), face-centered cubic (FCC), and spherical hollow cu-
bic (SHC) [189] . Comparatively, the FCC scaffold showed the lowest
stress concentration and highest compression properties. The ex-
perimental elastic modulus and compressive strength of the four
Co-Cr scaffolds ranged from 7.18 to 16.57 GPa and from 271.53 to
1279.52 MPa, respectively.
Among different L–PBF techniques, SLM has proved to be one
of the best methods for the fabrication of various types of ortho-
pedic scaffolds or augmentations. SLM is a multipurpose system
that allows for the fabrication of not only dense materials with a
very-well defined external shape, but also porous materials with a
designed pore shape, size, volume fraction, interconnectivity, and
occasionally gradient porosity [ 80 , 190 ]. SLM allows production of
fine and small porous Ti structures, with struts in the range of
10 0–20 0 μm. This enables the possibility of tailoring and optimiz-
ing the structural and mechanical properties of the scaffolds while
maintaining the required pore dimensions that allow for bone and
vessel ingrowth.
In a research study conducted in a load-bearing critical femoral
bone defect in rats, it was reported that SLM porous Ti scaffolds
209

A. Nour i, A. Ro hani Shirvan, Y. Li et al. Journal of Materials Science & Technology 94 (2021) 196–215
Fig. 11 . SEM macrographs of an SLM porous 316L stainless steel from: a) perpendicular angle; and b) oblique angle [78] . Reproduced with the permission from Elsevier.
with reduced strut size led to a reduction in stiffness of the struc-
ture and showed a positive effect on bone formation [191] . A 316L
stainless steel component with a gradient porous microstructure
was fabricated by SLM process through setting a gradient varia-
tion in scanning speed in each layer. Besides the density, the ten-
sile strength of the SLM 316L component was also dependent on
scan speed [173] . Using SLM, Pattanayak et al. [80] manufactured
cancellous bone-like biomimetic Ti scaffolds. An increase in com-
pressive strength from 35 MPa to 120 MPa was reported when the
porosity decreased from 75% to 55%. In another study, Shishkovsky
et al. [37] investigated the technical aspects of SLS in producing
porous load-bearing implants of NiTi and Ti alloys. Macroscopic
grooves or porous/rough surface on dental implants can not only
promote mechanical stabilization between implants and surround-
ing bone, but also provide better and faster bone apposition [158] .
Through a combined process, Tolochko et al. [157] fabricated dental
root implants with a compact core by SLM and a porous shell by
SLS, with a porosity range of 40–45% and the pore size of 10 0–20 0
μm. Capek and his co-workers [78] fabricated a porous 316L stain-
less steel scaffold with square-shaped and interconnected pores
by SLM and compared its microstructure, mechanical properties
and biocompatibility with those of dense (nonporous) 316L stain-
less steel prepared by SLM or casting followed by hot forging. The
fabricated porous structures exhibited relatively close mechanical
properties to those of cancellous bone. The macrostructure of the
porous 316L stainless steel with square-shaped pores is seen in
Fig. 11 . The images were acquired from different angles, show-
ing pores with a side length of approximately 750 μm and rough
cylindrical struts with a diameter of approximately 250 μm [78] .
Soro et al. [70] fabricated a porous Ti–25Ta alloy using SLM as a
promising and cost-effective load-bearing orthopedic implant. The
fabricated scaffold showed significantly reduced elastic modulus
ranging from 14 to 36 GPa (depending on the level of porosity)
which falls within the elastic modulus of cortical bone. These SLM
as-printed samples are shown in Fig. 12 . A noticeable layer of un-
melted powder particles is visible on the resulting non-treated sur-
face of the samples.
5.2. Porous polymeric scaffolds
Like metals, an ideal porous polymer scaffold should exhibit
ease of processability, biocompatibility, excellent mechanical prop-
erties, and preferably, biodegradability [192] . These properties can
be achieved with some biodegradable polymers. Biodegradable
polymers, such as PLA, polyglycolic acid (PGA), and poly(lactic-co-
glycolic acid) (PLGA), have attracted significant attention due to
their ease of fabrication and their relative high elasticity; however,
their degradation rate has to match the healing time of bone tis-
sue, should be neither too fast nor too slow (proximately 6 months
in vivo ) [193] . In other words, for bone tissue regeneration, the
scaffold should degrade over time to allow full regeneration of na-
tive tissue structure and its function. Rapid degradation rate leads
to the collaps of porous scaffold structure, and as a result, hinders
mass transfer and causes necrosis. In contrast, if the degradation
rate is too slow, tissue regeneration may be hampered by fibrotic
encapsulation and lack of host integration. Therefore, a suitable
degradation rate of scaffold is crucially important to achieve op-
timal bone tissue regeneration [ 194 , 195 ]. Generally, polymers and
organic scaffolds can be shaped into various suitable structural ar-
chitectures. However, the mechanical properties of the polymers
are usually lost, which often results in very low or no compres-
sive strength [163] . To strike a balance between good mechanical
properties and to achieve fully interconnected pores, PCL has been
shown to be a suitable candidate for porous polymeric scaffolds.
PCL also has a slower degradation rate than other polymers which
matches the rate of new bone formation and thus avoids possible
gaps in regeneration [196] .
Among L–PBF technologies, SLS is the only technique that is
used for the fabrication of porous polymer scaffolds necessary for
bone cell repair and regeneration. Williams et al. [65] designed
and fabricated PCL mandibular condyle scaffolds via SLS. The re-
sults confirmed a porous structure and sufficient mechanical prop-
erties of PCL scaffolds which made them suitable for bone tissue
engineering applications. Compressive modulus and yield strength
values were reported in the range of 52 to 67 MPa and 2.0 to 3.2
MPa, respectively. SLS-printed PCL scaffolds have also been used
to repair craniofacial bone defects. Smith et al. [197] designed a
promising scaffold for bone-tissue regeneration in craniomaxillofa-
cial surgery by the SLS method. PCL was used to create a ramus-
condyle unit scaffold f or application in a jaw joint reconstruction
in a Yucatan minipig animal model. Bone volumes and tissue min-
eral density at different time periods demonstrated significant new
bone growth interior and exterior to the scaffold. Thus, the SLS-
fabricated scaffold could support masticatory function as well as
both osseous and cartilage regeneration. In another study, a porous
PVA scaffold f or bone-tissue engineering was prepared by SLS. PVA
has a similar tensile strength to human articular cartilage there-
fore it is a promising polymer for the treatment of load-bearing
joint and craniofacial bone defects. The scaffold was fabricated un-
der a laser power of 8 W, scan speed of 600 mm/min, laser spot
diameter of 0.8 mm and layer thickness of 0.15 mm. The results
indicated that if the laser power was higher than 16 W, thermal
decomposition in the scaffolds would occur. The number of human
osteosarcoma MG-63 cells and the bridge between cells increased
with increasing cell culture time. Thus, SLS appears to be a promis-
210

A. Nour i, A. Ro hani Shirvan, Y. Li et al. Journal of Materials Science & Technology 94 (2021) 196–215
Fig. 12 . Images of dense and porous Ti–25Ta samples with different levels of porosity produced by SLM [70] . Reproduced with permission from Elsevier.
ing method in the fabrication of PVA scaffolds for bone-tissue en-
gineering [198] .
5.3. Porous composite scaffolds
In many clinical cases, porous composite scaffolds may prove
necessary for the reconstruction of bone defects. Common material
combinations are a synthetic polymer matrix with a bioceramic re-
inforcement and a synthetic/natural polymer matrix with a metal
reinforcement. Metal-ceramic-polymer hybrid materials have also
been proposed for the fabrication of load-bearing scaffolds [162] .
Recently, researchers have emphasized the use of composite
bone scaffolds, which combine the flexibility and toughness of
biodegradable organic matrixes (e.g. a polymeric matrix) and the
strength and stiffness of bioactive inorganic fillers (e.g. bioactive
ceramics) [199] . For orthopedic scaffolds designed for large size de-
fects, stiffness is a crucial parameter. Polymer-matrix scaffolds con-
taining bioactive ceramics can be manufactured to serve two pur-
poses: (a) making the scaffolds osteoconductive and (b) reinforcing
the scaffolds [36] . Ceramic materials have high stiffness. Therefore,
with increasing ceramic content in a polymer-ceramic composite,
the stiffness of the scaffold is increased. Nevertheless, obtaining
suitable mechanical properties for hard-tissue applications is diffi-
cult using porous polymer-ceramic composites. In particular, scaf-
folds based on HA or β-TCP are very stiff and brittle and may have
different viscoelastic properties from bone [200] .
SLS manufacturing of PCL-based scaffolds with natural/synthetic
polymers or bioceramic particles has proven to have great poten-
tial for bone-tissue applications. In a study, alginate and polyacry-
lamide (PAM) were used in the PCL scaffolds to improve their per-
formance. The mechanical analysis indicated that the elastic mod-
ulus increased from 6.99 MPa in pure PCL to 12.67 MPa in the bio-
composite, and the elongation at break of the samples increased
from 59% to 112.9%. In addition, cell seeding and in vitro culture
showed that the cell viability remained above 94% over 5 days. The
results showed the suitability of the fabricated PCL/alginate/PAM
scaffolds f or skelet al tissue repair [201] .
Eosoly et al. [202] used the SLS process to fabricate a biocom-
posite made of PCL with 30 wt.% HA in order to increase the cell
attachment, proliferation and differentiation. Based on the results,
the dimensions of the fabricated PCL/HA parts were strongly de-
pendent on the manufacturing direction and scan spacing parame-
ters, and they had significant effects on the accuracy and mechan-
ical properties. The compressive modulus and yield strength of the
scaffolds were between 0.6 and 2.3 and 0.1 and 0.6 MPa, respec-
tively. In another study by Doyle et al. [203] , orthopedic scaffold
materials fabricated from PCL/ β-TCP powders using SLS were in-
vestigated. Incorporating β-TCP particles is desirable to promote
osteoconductivity. The results indicated that with increasing ce-
ramic content, a significant increase in stiffness and a marked
reduction in strength were observed. The results highlighted the
influence of ceramic content on the mechanical properties and
degradation behavior of PCL/ β-TCP composites, and indicated that
these changes must be considered in the design of scaffolds for
critical-sized defects. In the same study, fabrication of PCL compos-
ites filled with different volume fractions of β-TCP (10%-30%) by
SLS for bone-tissue engineering was investigated. The optimal pro-
cessing parameters for each composition were developed by design
of experiments (DOE). This research provides valuable information
for fabricating tissue engineering scaffolds for load-bearing appli-
cations [204] .
Demineralized bone matrix (DBM) is an allograft bone ma-
trix with osteoinductive potential for bone scaffolds which is only
available in limited sizes. However, AM methods provide the op-
portunity to fabricate larger geometries while maintaining the ben-
efits of DBM. Ziaee et al. [205] fabricated scaffolds based on a
blend of DBM and PCL using SLS. The DBM/PCL blend was fused
in the SLS machine according to the optimal parameters and then
was strengthened through a heat treatment. The tensile strength
values indicated the vital role of post-processing heat treatment
in increasing the tensile strength of the DBM. Additionally, the fi-
nal part had 40% porosity with a density comparable to blocks of
DBM.
One of the main challenges in structural medical applications
is design of a biomimetic scaffold for osteochondral repair. To this
aim, a multilayer osteochondral scaffold based on PCL/HA micro-
spheres was fabricated by SLS. The scaffolds showed an excellent
biocompatibility and improved articular cartilage formation by ac-
celerating early subchondral bone regeneration [206] . Zhou et al.
[199] fabricated bio nanocomposite microspheres based on carbon-
ated hydroxyapatite (CHA) nanospheres within a poly (L-lactide)
(PLLA) matrix via SLS to produce bone-tissue-engineering scaffolds.
This platform allowed the use of small quantities of biomaterials
for the scaffold production. Porous scaffolds were successfully fab-
ricated from the PLLA microspheres and PLLA/CHA nanocompos-
ite microspheres. The results demonstrated the suitability of the
PLLA/CHA nanocomposite for the production of porous bone scaf-
folds via SLS. Using the same fabrication method, Huang et al.
[207] utilized coral HA/PLLA to fabricate porous bone scaffolds
with a rough surface. With increasing the proportion of coral HA
particles, the porosity, density, and hydrophilicity of the scaffolds
increased, while the compressive strength decreased. Moreover,
coral HA showed no negative impact on cell viability and prolif-
eration.
211

A. Nour i, A. Ro hani Shirvan, Y. Li et al. Journal of Materials Science & Technology 94 (2021) 196–215
Nanoclay is another filler for the preparation of PLA-based com-
posites for orthopedic applications. Clays are mineral materials
which are low cost and environmentally friendly. They are incorpo-
rated as a filler in the polymer matrix in order to improve the me-
chanical, thermal, and anticorrosion properties. Therefore, nanoclay
has wide usage in structural clinical applications including bone-
tissue engineering and bone cement [208] . Bai et al. [209] inves-
tigated the feasibility of processing of PLA and PLA/nanoclay com-
posite by a laser sintering method. It was shown that the powder
bed temperature for PLA/nanoclay was lower than that for pure
PLA. In addition, the PLA/nanoclay part had a better flexural mod-
ulus than pure PLA.
For creating a structural scaffold, a PVA/HA biocomposite was
produced using SLS and the feasibility of powder sintering and the
effects of process parameters on the final product were studied.
The obtained results showed that the sintering process did not af-
fect the chemical composition of the biocomposite and the laser-
sintered scaffold was bioactive in SBF [210] . Another biocompati-
ble polymer with high load-bearing capacity is PA. However, this
polymer shows a poor cell seeding behavior. Therefore, a combi-
nation of PA and a bioceramic like HA can provide a composite
scaffold with high compression strength, good biocompatibility and
cell seeding behavior. Through the SLS method, a PA/HA composite
was employed for the fabrication of load-bearing bone scaffolds.
The porous scaffolds were designed with different pore configura-
tions including cubic pores, spherical pores, shifted cubical pores,
and shifted spherical pores. According to the structural analysis,
the shifted cubic scaffold showed the minimum stress concentra-
tion, and thus, was selected as a suitable pore configuration. The
investigation proved that the SLS method offers the feasibility of
fabricating parts with complex geometries, good quality, high me-
chanical strength, and fine accuracy for bone regeneration appli-
cations [211] . In another study, PA/HA scaffolds with shifted cubic
pore configuration were fabricated using SLS in a vertical build di-
rection. The in vitro assessment of the fabricated scaffolds yielded
positive results for cell growth and the PA/HA (80:20) composition
exhibited maximum strengths of 24.3 MPa and 28.1 MPa during
tensile and compression tests, respectively [212] . Kamarajan et al.
[213] studied the suitability of a PA/HA composite orthopedic scaf-
fold. PA/HA (90:10) laser-sintered scaffolds were implanted in the
rat models. The obtained results showed that the PA/HA composite
scaffolds perf ormed consistently better under in vitro and in vivo
conditions in comparison with pure PA. Thus, it could more likely
be used in orthopedic applications.
6. Summary and future trends
The past few years have shown significant advances in AM tech-
nologies, leading to the production of biomaterials with structural
rigidity and high load-bearing capacity. As a subtype of AM, L–
PBF allows for rapid fabrication of complex geometrical and cus-
tomized implants made of various metallic and polymeric materi-
als. L–PBF is a leading technology in producing structural bioma-
terials for orthopedics, traumatology, craniofacial/maxillofacial and
dental applications, that offers ease of manufacturing and low ma-
terial wastage. The most available L–PBF methods are classified
into three groups: selective laser sintering (SLS); selective laser
melting (SLM); and direct metal laser sintering (DMLS), of which
SLS is mainly used in the processing of polymers and ceramics,
while SLM and DMLS are extensively utilized in the manufactur-
ing of metals and alloys.
The success of load-bearing implants fabricated through these
methods largely relies on their ability to emulate the mechanical
behavior of bone and to promote osseointegration for early fixa-
tion and long-term stability. Nevertheless, the mechanical proper-
ties, surface texture, dimensional accuracy, and manufacturing time
of the final products are considerably affected by the L–PBF pro-
cessing parameters including material properties, laser parameters
(e.g., laser power, scanning speed, hatch spacing, and layer thick-
ness), and powder characteristics. The current review has also pre-
sented some of the typical polymers and metals, along with their
properties, that are commonly used for L–PBF manufacturing of
load-bearing biomaterials. Although there is very little variance in
material compatibility among different L–PBF technologies, some
of them appear to support wider material selections likely due to
economic considerations. L–PBF has also proved to be a rapid and
cost-effective technology for manufacturing complex, accurate, and
mass-customized dental implants and load-bearing dental restora-
tions such as crowns, bridges, and dentures. By virtue of this tech-
nology, complete control over the microarchitecture and porosity
of metallic, polymeric, and composite scaffolds can be achieved.
The precisely designed pore configurations and orientations enable
tailoring and optimizing the structural properties of the implants
for increased biological/morphological fixation and bone ingrowth,
while maintaining a trade-off between porosity and mechanical
strength.
Despite the numerous advantages of L–PBF technology, it suf-
fers from some technical drawbacks. Only a small number of
metallic and polymeric materials can be processed by this method,
and the ease of fabrication of composite materials is an area of sig-
nificant challenge. Furthermore, the dimensional accuracy of the
designed porous scaffolds and the entrapment of powder within
them for orthopedic implants prove to be other challenging aspects
of L–PBF technology. Although 3D printers are becoming more af-
fordable with time, the costs of running and post-processing, ma-
terials, surface finishing, maintenance, and the need for skilled op-
erators must also be fully taken into account. It is also imperative
to have a thorough understanding of the anisotropy and hetero-
geneity in the microstructures and mechanical properties of L–PBF
printed metal parts that often occur due to complex cyclic thermal
histories consisting of directional heat extraction, repeated melting,
and rapid solidification.
Future developments in L–PBF technology are expected to
tackle the production costs, size limitations, quality consistency,
and to improve the mechanical properties and surface texture of
the final parts. Some intrinsic problems with the L–PBF produc-
tion of hard tissue implants and prosthetics can be further inves-
tigated and rectified, including dimensional imprecision associated
with shrinkage, warpage, waviness and surface roughness. Incorpo-
rating features such as hybrid design and functionally graded mi-
croarchitecture, are areas of investigation which can improve the
mechanical and biological properties of load-bearing implants. De-
velopments in laser technology, such as the introduction of the
femtosecond laser, could expand the use of a wider range of met-
als and alloys. Advanced equipment that is capable of printing
low melting point and reactive metals (e.g. Mg) in an inert atmo-
sphere is a research direction worth future investigation. In ad-
dition, more systematic research should be conducted to under-
stand the microstructural evolution during the L–PBF and to iden-
tify the powder properties that affect their spreading. As to amor-
phous polymer powders, strategies to enhance the degree of con-
solidation by increasing the flow and sintering rate would result
in the production of polymeric biomaterials with less porosity and
high strength. Nonetheless, it is worth emphasizing that no single
technology alone is able to meet all the structural and functional
requirements of metallic and polymeric biomaterials. Thus, when
printing 3D structures it is advisable to combine the best attributes
of L–PBF technologies with conventional methods and to examine
the accuracy, reproducibility, and safety of 3D-printed biomaterials
by means of long-term clinical studies. From another perspective,
a closer look should be given to AM as a novel technical oppor-
tunity and a new tool that could allow the manufacturers and re-
212

A. Nour i, A. Ro hani Shirvan, Y. Li et al. Journal of Materials Science & Technology 94 (2021) 196–215
searchers to be more creative in developing load-bearing biomate-
rials with new designs and properties in a shorter time and more
cost-effective way. However, at the present time it is not wise to
fully entrust AM with the task of fabricating all the required struc-
tural and functional biomaterials. New standards should be legis-
lated to carefully monitor the liability and clinical performance of
additively manufactured biomaterials before putting into operation
and to make sure that they match or surpass those of convention-
ally manufactured biomaterials.
Declaration of Competing Interests
The authors declare that they have no known competing finan-
cial interests or personal relationships that could appear to have
influenced the work reported in this paper.
Acknowledgments
The authors acknowledge the financial support for this research
by the Australian Research Council (ARC) through the discovery
grant DP170102557 . YL is also supported through an ARC Future
Fellowship ( FT160100252 ). The authors are thankful to the Bonash
Medical and Adeiss centre at Western University for providing the
images of craniomaxillofacial biomaterials. The author A. Nouri
would also like to express his gratitude to Prof. Nick Birbilis for
the images of SLM-fabricated acetabular hip cup. The authors also
gratefully acknowledge the KLS Martin Group for the image of
biodegradable patient-specific cranial implant.
References
[1] A. Dawood , B.M. Marti , V. Sauret-Jackson , A. Darwood , Brit. Dent. J. 219 (2015)
521–529 .
[2] C.W. Hull , Apparatus for Production of Three-Dimensional Objects by Stere-
olithography, United States Patent, Appl, 1984 No 638905, Filed .
[3] F. Auricchio , S. Marconi , EFORT Open Rev. 1 (2016) 121–127 .
[4] , Global Additive Manufacturing with Metal Powders Market 2021-2025 Covid
19 Impact on Top Countries Data Industry Growth Analysis, Segmentation,
Size, Share, Trend, Future Demand and Leading Players Updates by Forecast,
2021 https://www.marketwatch.com/press-release .
[5] B. Dutta , F.H .S . Froes , in: Titanium Powder Metallurgy, Elsevier, 2015,
pp. 447–468 .
[6] A .A . Zadpoor , Int. J. Mol. Sci. 18 (2017) 1607 .
[7] P. Ahangar , M.E. Cooke , M.H. Web er , D.H. Rosenzweig , Appl. Sci. 9 (2019)
1713 .
[8] A. Rohani Shirvan , A. Bashari , N. Hemmatinejad , Eur. Polym. J. 119 (2019)
541–550 .
[9] M. Averyanova , P. Bertrand , B. Verquin , Virtual Phys. Prototyp. 6 (2011)
179–185 .
[10] T. Nakano , T. Ishimoto , KONA Powder Part. J. 32 (2015) 75–84 .
[11] R. Galante , C.G. Figueiredo-Pina , A.P. Serro , Dent. Mater.
35 (2019) 825–846 .
[12] S. Derakhshanfar , R. Mbeleck , K. Xu , X. Zhang , W. Zhong , M. Xing , Bioact.
Mater. 3 (2018) 144–156 .
[13] A. Nouri , P.D . Hodgson , C. Wen , in: Biomimetics Learning From Nature, InTech,
2010, pp. 415–450 .
[14] A. Nouri , Novel Metal Structures Through Powder Metallurgy for Biomedical
Applications, Deakin University, 2008 Ph. D. Thesis .
[15] T. Sridhar , S. Vinodhini , U. Kamachi Mudali , B. Venkatachalapathy ,
K. Ravichandran , Mater. Tech. 31 (2016) 705–718 .
[16] M. Geetha , A.K. Singh , R. Asokamani
, A.K. Gogia , Prog. Mater. Sci. 54 (2009)
397–425 .
[17] L. Bai , C. Gong , X. Chen , Y. Sun , J. Zhang , L. Cai , S. Zhu , S.Q. Xie , Metals 9
(2019) 100 4 .
[18] M.J. Galba , T. Reischle , in: Additive Manufacturing, CRC Press, Boca Raton,
2015, pp. 97–142 .
[19] A. Nouri , C. We n , in: Structural Biomaterials: Properties, Characteristics, and
Selection, Elsevier, 2021, pp. 127–156 .
[20] A . Rohani Shirvan , A . Nouri , C. Wen , in: Structural Biomaterials: Properties,
Characteristics, and Selection, Elsevier, 2021, pp.
395–439 .
[21] L.J. Tan , W. Zhu , K. Zhou , Adv. Func. Mater. 30 (2020) 2003062 .
[22] B. Valmik , K. Prakash , P. Vinaykumar , K. Shreyans , G. Kiran , S. Rajkumar , A
Review on Powder Bed Fusion Technology of Metal Additive Manufacturing,
4th International Conference and Exhibition on Additive Manufacturing Tech-
nologies, 2014 4th International conference and exhibition on Additive Man-
ufacturing Technologies-AM-2014, Septeber .
[23] C. Han , Q. Fang , Y. Shi , S.B. Tor , C.K. Chua , K. Zhou , Adv. Mater. 32 (2020)
1903855 .
[24] C.Y. Yap , C.K. Chua , Z.L. Dong , Z.H. Liu , D.Q. Zhang , L.E. Loh , S.L. Sing , Appl.
Phys. Rev. 2 (2015) 041101 .
[25] A. Sola , A. Nouri , J. Adv. Manuf. Proc. 1 (2019) e10021 .
[26] D.L. Bourell , D.K. Leigh , Procedia CIRP 94 (2020) 144–148 .
[27] A . Nouri , A . Sola , in: Metallic Biomaterials Processing and Medical Device
Manufacturing, Elsevier, 2020, pp. 271–314 .
[28] J.P. Kruth , P. Mercelis , J.V. Vaerenbergh , L. Froyen , M. Rombouts ,Rapid Proto-
typing J. 11 (2005) 26–36
.
[29] J.P. Kruth, B. Vandenbroucke, J. Van Vaerenbergh, P. Mercelis, Benchmark-
ing of different SLS/SLM processes as rapid manufacturing techniques. Pro-
ceedings of the International Conference Polymers & Moulds Innovations PMI
2005.
[30] A. Khorasani , I. Gibson , J.K. Veetil , A.H. Ghasemi , Int. J. Adv. Manuf. Tech . 108
(2020) 191–209 .
[31] H. Zhang , S. LeBlanc , in: Additive Manufacturing of High-performance Metals
and Alloys: Modeling and Optimization, InTech, 2018, pp. 1–44 .
[32] C.R. Deckard , Method and Apparatus for Producing Parts by Selective Sinter-
ing, 19 89 Google PatentsUS4863538A .
[33] W. Ruban , V. Vijayakumar , P. Dhanabal , T. Pridhar , Int. J. Rapid Manuf. 4
(2014) 148–164 .
[34] F. Fina , A. Goyanes , S. Gaisford , A.W. Basit , Int. J. Pharmaceut. 529 (2017)
285–293 .
[35] J.J. Beaman , J.W. Barlow , D.L. Bourell , R.H. Crawford , H.L. Marcus , K.P. McAlea ,
in: Solid Freeform Fabrication: A New Direction in Manufacturing, Springer,
1997, pp. 245–278 .
[36] A. Mazzoli , Med. Biol. Eng. Comput. 51 (2013) 245–256 .
[37] I. Shishkovsky , L. Volova , M. Kuznetsov , Y.G. Morozov , I. Parkin , J. Mater.
Chem. 18 (2008) 1309–1317 .
[38] H. Yaagoubi , H. Abouchadi , M.T. Janan , in: Review on the Modeling of the
Laser Sintering Process for Polyamide 12. E3S Web of Conferences, EDP Sci-
ences, 2021, p. 0 0 0 06 .
[39] S.C. Ligon , R. Liska , J. Stampfl, M. Gurr , R. Mülhaupt , Chem. Rev. 117 (2017)
10212–10290 .
[40] D. Ibrahim , T. L. Broilo , C. Heitz , M.G. de Oliveira , H.W. de Oliveira , S.M.W. No-
bre , J.H.G. dos Santos Filho , D.N. Silva , J. Cranio Maxill. Surg. 37 (20 09)
167–173 .
[41] T.
Wohlers , T. Gornet , History of Additive Manufacturing, 2011 Wohlers Re-
port: Additive Manufacturing and 3D Printing State of the Industry Annual
Worldwide Progress Report .
[42] O. Neikov , in: Handbook of Non-Ferrous Metal Powders, Elsevier, 2019,
pp. 373–399 .
[43] D. Gu , Y.C. Hagedorn , W. Meiners , G. Meng , R.J.S. Batista , K. Wissenbach ,
R. Poprawe , Acta Mater. 60 (2012) 3849–3860 .
[44] T. Moritz , S. Maleksaeedi , in: Additive Manufacturing: Materials, Processes,
Quantifications and Applications, Butterworth-Heinemann, 2018, pp. 105–161 .
[45] F. Tamimi , H. Hirayama , Digital Restorative Dentistry, Springer,
2019 .
[46] C.Y. Liaw , M. Guvendiren , Biofabrication 9 (2017) 024102 .
[47] B. Bhushan , M. Caspers, Microsyst. Technol. 23 (2017) 1117– 112 4 .
[48] D. Dzhendov , T. Dikova , J. IMAB 22 (2016) 1414– 1417 .
[49] C. Gao , C. Wang , H. Jin , Z. Wang , Z. Li , C. Shi , Y. Leng , F. Yan g , H. Liu , J. Wang ,
RSC Adv. 8 (2018) 25210–25227 .
[50] R. Singh , S. Singh , M. Hashmi , Reference Module in Materials Science and
Materials Engineering in: S. Hashmi (Ed.), Elsevier, 2016 .
[51] A. Gratton , in: Proceedings of The National Conference On Undergraduate Re-
search (NCUR), Weber State University, Ogden Utah, 2012 .
[52] A. Jardini , M. Larosa , A. Kaasi , P. Kharmandayan , Reference Module in Materi-
als Science and Materials Engineering in: S. Hashmi (Ed.), Elsevier, 2017 .
[53] F. Mangano , L. Chambrone , R. Van Noort , C. Miller , P. Hatton , C. Mangano , Int.
J. Biomat. (2014) 461534 2014 .
[54] T. Duda , L.V. Raghavan , AI & Society 33 (2018) 241–252 .
[55] D. Martinez-Marquez , M. Jokymaityte , A. Mirnajafizadeh
, C.P. Carty , D. Lloyd ,
R.A. Stewart , Materials 12 (2019) 3110 .
[56] D. Przestacki , J. Sieniawski , G. Stambolov , P. Lisiak , Arch. Mech. Technol.
Mater. 35 (2015) 23–32 .
[57] E. Provaggi , J.J. Leong , D.M. Kalaskar , Proc. Inst. Mech. Eng. H 231 (2017)
471–486 .
[58] L. Yang , K. Hsu , B. Baughman , D. Godfrey , F. Medina , M. Menon , S. Wiener ,
Additive Manufacturing of Metals: The Technology, Materials, Design and Pro-
duction, Springer, 2017 .
[59] M. Schmid , A. Amado , K. Wegener , AIP Conference proceedings, AIP Publish-
ing LLC, 2015 .
[60] Y. Shi , Z. Li , H. Sun , S. Huang , F. Zeng , J. Mater , Des. Appl. 218 (2004) 247–252 .
[61] G. Barbosa-Canovas , Food Powders, Kluwer Academic Press, New York, 2005 .
[62] W. Zhu , C. Yan , Y. Shi , S. Wen , C. Han , C. Cai , J. Liu , Y. Shi , Rapid Prototyp. J.
22 (2016) 621–629 .
[63] J. Schmidt , M.. Sachs , S Fa nsel ow , M. Zhao , S. Romeis , D. Drummer , K.E. Wirth ,
W. Peukert , Chem. Eng. Sci. 156 (2016) 1–10 .
[64] J.P. Kruth , G. Levy , F. Klocke , T.H.C. Childs , CIRP Ann. 56 (2007) 730–759 .
[65] J.M. Williams , A. Adewunmi , R.M. Schek , C.L. Flanagan , P. H . Krebsbach ,
S.E. Feinberg , S.J. Hollister , S. Das , Biomaterials 26 (2005) 4 817–4 827 .
[66] J.P.Kruth B.Vandenbroucke , Rapid Prototyping J. 13 (20 07) 196–203 .
[67] A Nouri , C Wen , in: Structural Biomaterials: Properties, Characteristics, and
Selection, Elsevier, 2021, pp. 67–101 .
[68] W.H. Lee , Y. Zhang , J. Zhang ,
Powder Technol. 315 (2017) 300–308 .
[69] , Met. Powder Report 69 (2014) 24–27 .
213

A. Nour i, A. Ro hani Shirvan, Y. Li et al. Journal of Materials Science & Technology 94 (2021) 196–215
[70] N. Soro , H. Attar , E. Brodie , M. Veidt , A. Molotnikov , M.S. Dargusch , J. Mech.
Behav. Biomed. Mater. 97 (2019) 149–158 .
[71] H. Irrinki , S.D. Nath , A .A . Akilan , G. Gupta , S.V. Atre , Additive Manufacturing
Processes, ASM Handbook in: D. Bourell, W. Frazier, H. Kuhn, M. Seifi, (Eds.),
2020, p. 24 .
[72] B. Vandenbroucke , J.P. Kruth , in: Selective Laser Melting of Biocompat-
ible Metals for Rapid Manufacturing of Medical Parts, Proceedings of
Solid Freeform Fabrication Symposium, University of Texas at Austin, 2006,
pp. 148–159 .
[73] S. Das , M. Wohl ert , J.J. Beaman , D.L. Bourell , JOM 50 (1998) 17–20 .
[74] A. Adeyemi , E.T. Akinlabi , R.M. Mahamood , Mater. Today: Proceed 5 (2018)
18510 –18517 .
[75] J. Suryawanshi , K. Prashanth , U. Ramamurty , Mater. Sci. Eng. A 696 (2017)
113–121 .
[76] Y. Shang , Y. Yuan , D. Li , Y. Li , J. Chen , Int. J. Adv. Manuf. Technol. 92 (2017)
4379–4385 .
[77] P. Guo , B. Zou , C. Huang , H. Gao , J. Mater. Proc. Technol. 240 (2017) 12–22 .
[78] J ˇ
Capek , M Machová, M Fousová, J Kubásek , D Vojt
ˇ
ech , J Fojt , E. Jablonská,
J. Lipov , T. Ruml , Mater. Sci. Eng. C 69 (2016) 631–639 .
[79] J. Black , Biological Performance of Materials: Fundamentals of Biocompatibil-
ity, Crc Press, 2005 .
[80] D.K. Pattanayak , A. Fukuda , T. Matsushita , M. Take moto , S. Fujibayashi ,
K. Sasaki , N. Nishida , T. Nakamura , T. Kok ubo , Acta Biomater 7 (2011)
1398–1406 .
[81] S. Ozan , K. Munir , A. Biesiekierski , R. Ipek , Y. Li , C. Wen ,
in: Biomaterials Sci-
ence: An Introduction to Materials in Medicine, 4th Ed., Academic Pres, 2020,
pp. 229–247 .
[82] Q. Wang , C. Han , T. Choma , Q. Wei , C. Yan , B. Song , Y. Shi , Mater. Des. 126
(2017) 268–277 .
[83] D. Zhao , C. Han , J. Li , J. Liu , Q. Wei , Mater. Sci. Eng. C 111 (2020) 110784 .
[84] D Zhao , C Han , Y Li , J Li , K Zhou , Q Wei , J. Liu , Y. Shi , J. Alloy. Compd. 804
(2019) 288–298
.
[85] C Han , Y Li , Q Wang , D Cai , Q We i , L Yan g , S. Wen , J. Liu , Y. Shi , Mater. Des.
141 (2018) 256–266 .
[86] NS Moghaddam , S Saedi , A Amerinatanzi , A Hinojos , A Ramazani , J Kundin ,
M.J. Mills , H. Karaca , M. Elahinia, Sci. Rep. 9 (2019) 1–11 .
[87] M. Elahinia , N.S. Moghaddam , M.T. Andani , A. Amerinatanzi , B.A. Bimber ,
R.F. Hamilton , Prog. Mater. Sci. 83 (2016) 630–663 .
[88] C.C. Ng , M. Savalani , H.C. Man , I. Gibson , Virtual Phys. Prototyp. 5 (2010)
13–19 .
[89] C.C. Ng , M. Savalani , H.C. Man , Rapid Prototyp. J. 17 (2011) 479–490 .
[90] H. Kabir , K. Munir , C. Wen , Y. Li , Bioact. Mater. 6 (2021) 836–879 .
[91] R. Wau thl e , J. Van Der Stok , S.A. Yava ri , J. Van Humbeeck , J.P. Kruth , A .A . Zad-
poor , H. Weinans , M. Mulier , J. Schrooten , Acta Biomater 14 (2015) 217–225 .
[92] J. Henckel , T.J. Holme , W. Radford , J.A .
Skinner , A .J. Hart , J. Amer. Acad. Or-
thop. Surg. 26 (2018) e342–e348 .
[93] T. Sözen , L. Özı¸s ık , N.Ç. Ba ¸s aran , Eur. J. Rheumatol. 4 (2017) 46–56 .
[94] S.H. Masood , in: Comprehensive Materials Processing, Elsevier, Oxford, 2014,
pp. 1–2 .
[95] P. Bartolo, J.P. Kruth, J. Silva, G. Levy, A. Malshe, K. Rajurkar, M. Mitsuishi, J.
Ciurana, M. Leu, CIRP Ann. 61 (2012) 635-655.
[96] P. Bártolo , C. Chua , H. Almeida , S. Chou , A. Lim , Virtual Phys. Prototyp. 4
(2009) 203–216 .
[97] L. Polo-Corrales , M. Latorre-Esteves , J.E. Ramirez-Vick ,J. Nanosci. Nanotech-
nol. 14 (2014) 15–56 .
[98] J.M. Broyles , N.B. Abt , S.M. Shridharani , B. Bojovic , E.D. Rodriguez , A.H. Doraf-
shar , Plast. Reconstr. Surg. 134 (2014) 760–769 .
[99] E.L. Nyberg , A.L. Farris , B.P. Hung , M. Dias , J.R. Garcia , A.H. Dorafshar ,
W.L. Grayson , Ann. Biomed. Eng. 45 (2017) 45–57 .
[100] S.J. Hollister , C. Lin , E. Saito , C. Lin , R. Schek , J. Taboas , J.M. Williams , B Partee ,
C.L. Flanagan , A. Diggs , E.N. Wilke , G.H. Van Lenthe , R. Müller , T. Wirtz , S. Das ,
S.E. Feinberg , P.H . Krebsbach , Orthod. Craniofac. Res. 8 (2005) 162–173 .
[101] S.P. Narra , P.N. Mittwede , S.D. Wolf , K.L. Urish , Orthop. Clin. 50 (2019) 13–20 .
[102] W.M. Mihalko , in: 3D Printing in Orthopaedic Surgery, Elsevier, 2019,
pp. 49–53 .
[103] R. Velu , T. Calais , A. Jayakumar , F. Raspall , Materials 13 (2020) 92 .
[104] M. Javaid , A. Haleem , J. Clin. Orthop. Trauma 9 (2018) 202–206 .
[105] A .L. Jardini , M.A . Larosa , C.A . de Carvalho Zavaglia , L.F. Bernardes , C.S. Lam-
bert , P. Kharmandayan , D. Calderoni , R.M. Filho ,Virtual Phys. Prototyp. 9
(2014) 115–125 .
[106] D Wan g , Y Wang , S Wu , H Lin , Y Ya ng , S Fan ,
C. Gu , Materials 10 (2017) 35 .
[107] A .R. Shirvan , A . Nouri , in: Advances in Functional and Protective Te xtiles, El-
sevier, 2020, pp. 291–333 .
[108] A. Nouri , X. Chen , P.D. Hodgson , C. Wen , Adv. Mater. Res. 15 (2007) 71–76 .
[109] K.A. Nazari , A. Nouri , T. Hilditch , Mater. Des. 88 (2015) 1164–1174 .
[110] A.M. Shah , H. Jung , S. Skirboll , Neurosurg. Focus 36 (2014) E19 .
[111 ] I.V. Panayotov , V. Orti , F. Cuisinier , J. Ya ch ou h ,J. Mater. Sci. Mater. Med. 27
(2016)
118 .
[112] K. Tan , C. Chua , K. Leong , C. Cheah , P. Cheang , M.A. Bakar , S.W. Cha , Biomater.
24 (2003) 3115 –3123 .
[113] J.T. Rimell , P.M. Marquis , J. Biomed. Mater. Res. 53 (20 0 0) 414–420 .
[114] A. Aksu , F. Reinauer , T. Wolfra m , Implant Production Method Using Additive
Selective Laser Sintering, and Implant, 2019 Google Patents .
[115] Z. Li , M. Cheng , J. Zhao , J. Hou , Selected Region Laser Sintering and Molding
Method of Ultrahigh Molecular Weight Polyethylene, 2015 China Patent .
[116] S. Changhui , H. Aibing , Y. Yongqiang , W. Di , Y. Jia-kuo , Int. J. Adv. Manuf. Te ch-
nol. 85 (2016) 1217–1226 .
[117] H. Ren , A. Li , B. Liu , Y. Dong , Y. Tian , D. Qiu , Sci. Rep. 7 (2017) 1–10 .
[118] A. Bistolfi, R. Ferracini , C. Albanese , E. Vernè, M. Miola , Materials 12 (2019)
4002 .
[119] R. Velu , S. Singamneni , J. Mater. Res. 29 (2014) 1883–1892 .
[120] R. Velu , S. Singamneni , in: International Solid Freeform Fabrication Sympo-
sium-An Additive Manufacturing Conference,
Austin TX, 2015, pp. 1738–1749 .
[121] M. van Dijk , T.H. Smit , E.H. Burger , P.I. Wuisman , Spine 27 (2002) 2706–2714 .
[122] M. Kanayama , B.W. Cunningham , C.J. Haggerty , K. Abumi , K. Kaneda ,
P.C. McAfee , J. Neurosurg. 93 (20 0 0) 259–265 .
[123] P.J. Rao , M.H. Pelletier , W.R. Wa lsh , R.J. Mobbs , Orthop. Surg. 6 (2014) 81–89 .
[124] C.Y. Lin , T. Wirtz , F. LaMarca , S.J. Hollister , J. Biomed. Mater. Res. A 83 (2007)
272–279 .
[125] J.L. Burnard , W.C. Parr , W.J.
Choy , W.R . Wal sh , R.J. Mobbs ,Eur. Spine J. 29
(2020) 1248–1260 .
[126] C Chen , Y Hao , X Bai , J Ni , S-M Chung , F Liu , I.S. Lee , Mater. Des. 17 5 (2019)
107824 .
[127] S. Vadapalli , K. Sairyo , V.K. Goel , M. Robon , A. Biyani , A. Khandha , N.A. Ebra-
heim , Spine 31 (2006) E992–E998 .
[128] C. Basgul , T. Yu , D.W. MacDonald , R. Siskey , M. Marcolongo , S.M. Kurtz , J.
Mater. Res. 33 (2018) 2040–2051 .
[129] M. Tilton , G.S. Lewis , G.P. Manogharan , in: Orthopedic Biomaterials:
Progress in Biology, Manufacturing, and Industry Perspectives, Springer, 2018,
pp. 21–55 .
[130] M. Cronskär , L.E. Rännar , M. Bäckström , K. Nilsson , B. Samuelsson , J. Mech.
Des. 13 7 (2015) 111418 .
[131] M. Cronskär , On Customization of Orthopedic Implants-From Design and Ad-
ditive Manufacturing to Implementation, Mid Sweden University, Kopierin-
gen, 2014 Ph.D. Thesis .
[132] G.S. Lewis , C.T. Caroom , H. Wee , D. Jurgensmeier , S.D. Rothermel ,
M.A. Bramer , J.S. Reid , J. Orthop. Trauma 29 (2015) e364 .
[133] N. Fiz ,
D. Delgado , X. Sánchez , P. Sánchez , A.M. Bilbao , J. Oraa , M. Sánchez ,
Ann. Transl. Med. 5 (2017) 400 .
[134] S. Singare , L. Yaxi on g , L. Dichen , L. Bingheng , H. Sanhu , L. Gang , Rapid Proto-
typ. J. 12 (2006) 206–213 .
[135] P. Stoor , A. Suomalainen , C. Lindqvist , K. Mesimäki , D. Danielsson , A. Wes ter-
mark , R.K. Kontio , J. Craniomaxillofac. Surg. 42 (2014) 1644–1649 .
[136] A. Bhargav , V. Sanjairaj , V. Rosa , L.W. Feng , J.F. Yh ,J. Biomed. Mater. Res.
B
106 (2018) 2058–2064 .
[137] R.F. Gregolin , F.M. Barbosa , C.A. de Camargo Zavaglia , R.C. Tokimat su ,in: Pro-
ceedings of 22
nd
International Congress of Mechanical Engineering, 2013 .
[138] S. Berretta , K. Evans , O. Ghita , Mater. Des. 139 (2018) 141–152 .
[139] C.A . Pimentel , R.A . Rezende , M.F. Oliveira , J.P Lopes da Silva , M.V.L. Fook , Ad-
ditive Manufacturing by Selective Laser Sintering:
Poly(ether ether ketone)
Processing Development, 38
th National Meeting of Production Engineering
Brazil, 2018 Aberpo .
[140] M Roskies , JO Jordan , D Fa ng , M-N Abdallah , MP Hier , A Mlynarek , F. Tamimi ,
S.D. Tran , J. Biomater. Appl. 31 (2016) 132–139 .
[141] P. Fiaschi , M. Pavanello , A. Imperato
, V. Dallolio , A. Accogli , V. Capra , A. Con-
sales , A. Cama , G. Piatelli , J. Neurosurg. Pediatr. 17 (2016) 705–710 .
[142] C. Schiller , M. Epple , Biomater. 24 (2003) 2037–2043 .
[143] C. Gayer , J. Ritter , M. Bullemer , S. Grom , L. Jauer , W. Meiners , A. Pfister ,
F. Reinauer , M. Vucak , K. Wissenbach , H. Fischer , R. Poprawe , J.H. Schleifen-
baum , Mater. Sci. Eng. C 101 (2019) 660–673 .
[144] H. Bafti , A. Habibolahzadeh , Mater. Des. 31 (2010) 4122–4129 .
[145] R. Van Noort , Dent. Mater. 28 (2012) 3–12 .
[146] R. Zandparsa , Dent. Clin. 58 (2014) 135–158 .
[147] S. Prasad , N.A. Kader , G. Sujatha , T. Raj , S. Patil , 3D Printing in Dentistry, 2018
Future Medicine .
[148] M. Javaid , A. Haleem , J. Oral. Biol. Craniofac. Res. 9 (2019) 179–185 .
[149] R. Williams , R. Bibb , D. Eggbeer , J. Collis , J. Prosthet. Dent. 96 (2006) 96–99 .
[150] J.R. Strub , E.D. Rekow , S. Witkowski , J. Am. Dent. Assoc. 137 (20 06)
1289–1296 .
[151] L.
Wu , H. Zhu , X. Gai , Y. Wang , J. Prosthet. Dent. 111 (2014) 51–55 .
[152] S.L. Chang , C.H. Lo , C.P. Jiang , D.J. Juan , Int. J. Pharma Med. Biol. Sci. 4 (2015)
184–187 .
[153] J. Chen , R. Ahmad , H. Suenaga , W. Li , K. Sasaki , M. Swain , Q. Li , PloS One 10
(2015) e0132552 .
[154] S.J. Lee , I.Y. Jung , C.Y. Lee , S.Y. Choi , K.Y. Kum , Dent. Traumatol. 17 (2001)
114–119 .
[155] J. Chen , Z. Zhang , X. Chen ,
C. Zhang , G. Zhang , Z. Xu , J. Prosthet. Dent. 112
(2014) 1088–1095 .
[156] S. Tunchel , A. Blay , R. Kolerman , E. Mijiritsky , J.A. Shibli , Int. J. Dent. (2016)
8590971 2016 .
[157] N. Tol oc hko , V. Savich , T. Laoui , L. Froyen , G. Onofrio , E. Signorelli , V.I. Titov ,
Proc. Instn. Mech. Engrs. L J. Mater. Des. Appl. 216 (2002) 267–270 .
[158] T. Traini , C. Mangano , R. Sammons , F. Mangano , A. Macchi , A. Piattelli , Dent.
Mater. 24 (2008) 1525–1533 .
[159] P.
Davies , M. Kearns , in: Medical Device Materials IV: Proceedings from the
Materials & Processes for Medical Devices Conference 2007, ASM Interna-
tional, 20 08, p. 42 .
[160] P.H. Lee , H. Chung , S.W. Lee , J Yoo , J. Ko , Int. Manuf. Sci. Eng. Conf. (2014)
V0 01T 0 04A045 .
[161] E.O. Olakanmi , R. Cochrane , K. Dalgarno , Prog. Mater. Sci. 74 (2015) 401–477 .
[162] K. Alvarez , H. Nakajima , Materials 2 (2009) 790–832 .
214

A. Nour i, A. Ro hani Shirvan, Y. Li et al. Journal of Materials Science & Technology 94 (2021) 196–215
[163] M. Pilia , T. Guda , M. Appleford , BioMed. Res. Int. (2013) 458253 2013 .
[164] A. Nouri , P.D . Hodgson , C.E. Wen , Acta Biomater 6 (2010) 1630–1639 .
[165] K.A. Nazari , T. Hilditch , M.S. Dargusch , A. Nouri , J. Mech. Behav. Biomed.
Mater. 63 (2016) 157–163 .
[166] A. Nouri , in: Metallic Foam Bone: Processing, Modification and Characteriza-
tion, Elsevier, 2017, pp. 131–160 .
[167] T. Majumdar , N. Eisenstein , J.E. Frith , S.C. Cox , N. Birbilis , Adv. Eng. Mater. 20
(2018) 180 0172 .
[168] M. Wong , J.
Eulenberger , R. Schenk , E. Hunziker , J. Biomed. Mater. Res. 29
(1995) 1567–1575 .
[169] L. Carlsson , T. Röstlund , B. Albrektsson , T. Albrektsson , Int. J. Oral Maxillofac.
Implants 3 (1988) 21–24 .
[170] K.A. Nazari , A. Nouri , T. Hilditch , Mater. Lett. 140 (2015) 55–58 .
[171] A. El-Hajje , E.C. Kolos , J.K. Wang , S. Maleksaeedi , Z. He , F.E. Wiria , C. Choong ,
A.J. Ruys , J. Mater. Sci. Mater. Med. 25 (2014) 2471–2480 .
[172] D. Melancon , Z. Bagheri , R. Johnston , L. Liu ,
M. Tanzer , D. Pasini , Acta Bio-
mater. 63 (2017) 350–368 .
[173] R. Li , J. Liu , Y. Shi , M. Du , Z. Xie , J. Mater. Eng. Perform. 19 (2010) 666–671 .
[174] A. Sola , D. Bellucci , V. Cannillo , Biotechnol. Adv. 34 (2016) 504–531 .
[175] L. Yan g , C. Han , H. Wu , L. Hao , Q. Wei , C. Yan , Y. Shi , J. Mech. Behav. Biomed.
Mater. 109 (2020) 103843 .
[176] C. Han , Y. Li , Q. Wa ng , S. We n , Q. Wei , C. Yan , L. Hao , J. Liu , Y. Shi , J. Mech.
Behav. Biomed. Mater. 80 (2018) 119–127 .
[177] D.D. Lima , S.A. Mantri , C.V. Mikler , R. Contieri , C.J. Yannetta , K.N. Campo ,
E.S. Lopes , M.J. Styles , T. Borkar , R. Caram , R. Banerjee , Mater. Des. 13 0 (2017)
8–15 .
[178] B. Derby , Science 338 (2012) 921–926 .
[179] B. Levine , Adv. Eng. Mater. 10 (20 08) 788–792 .
[180] J. Parthasarathy , Ann. Maxillofac. Surg. 4 (2014) 9–18 .
[181] D.C. Dunand , Adv. Eng. Mater. 6 (2004) 369–376
.
[182] J. Banhart , Prog. Mater. Sci. 46 (2001) 559–632 .
[183] L. Yuan , S. Ding , C. Wen , Bioact. Mater. 4 (2019) 56–70 .
[184] L.J. Gibson , M.F. Ashby , Cellular Solids: Structure and Properties, 2nd Ed.,
Cambridge University Press, 1997 .
[185] M.F. Ashby , A. Evans , N.A. Fleck , L.J. Gibson , J.W. Hutchinson , H.N.G. Wadley ,
Metal Foams: A Desgin Guide, Butterworth-Heinemann, Boston, 20 0 0 .
[186] C. Pan , Y. Han , J. Lu , Appl. Sci. 10 (2020) 6374 .
[187] H. Chen , Q. Han , C.
Wang , Y. Liu , B. Chen , J. Wang , Front. Bioeng. Biotechnol.
8 (2020) 609 .
[188] C. Han , Y. Yao , X. Cheng , J. Luo , P. Luo , Q. Wang , F. Yang , Q. Wei , Z. Zhang ,
Biomacromolecules 18 (2017) 3776-3378 .
[189] C. Han , C. Ya n , S. Wen , T. Xu , S. Li , J. Liu , Q. Wei , Y. Shi ,Rapid Prototyp. J. 23
(2017) 16–27 .
[190] C. Yan , L. Hao , A. Hussein , P. You ng , J. Mech. Behav. Biomed. Mater. 51 (2015)
61–73 .
[191] J. Van der Stok , O.P. Van der Jagt , S. Amin Yavari , M.F. De Haas , J.H. Waars-
ing , H. Jahr , E.M.M. Van Lieshout , P. Patka , J.A.N Verhaar , A .A . Zadpoor ,
H. Weinans , J. Orthop. Res. 31 (2013) 792–799 .
[192] R.S. Ambekar , B. Kandasubramanian , Ind. Eng. Chem. Res. 58 (2019)
6163–6194 .
[193] H. Wang , Y. Li , Y. Zuo , J. Li , S. Ma , L. Cheng , Biomaterials 28 (2007)
3338–3348 .
[194] S.H. Park , E.S. Gil , H. Shi , H.J. Kim , K. Lee , D.L. Kaplan , Biomaterials 31 (2010)
6162–6172 .
[195] Y. Shi , T. Pan , W. Zhu , C. Ya n , Z. Xia , J. Mech. Behav.Biomed. Mater. 10 4 (2020)
10366 4 .
[196] I. Gibson , M.M. Savalani , X.L. Christopher , R. Olkowski , A.K. Ekaputra , K.C. Tan ,
D.W. Hutmacher , Tsinghua Sci. Technol. 14 (2009) 13–19 .
[197] M. Smith , C. Flanagan , J. Kemppainen , J. Sack ,
H. Chung , S. Das , S.J. Hollister ,
S.E. Feinberg , Int. J. Med. Robotics Comput. Assist. Surg. 3 (2007) 207–216 .
[198] C. Shuai , Z. Mao , H. Lu , Y. Nie , H. Hu , S. Peng , Biofabrication 5 (2013) 015014 .
[199] W.Y. Zhou , S.H. Lee , M. Wang , W.L. Cheung , W. Y. Ip , J. Mater. Sci. Mater. Med.
19 (2008) 2535–2540 .
[200] E. Giesen , N. Lamerigts , N. Verdonschot , P. Buma , B. Schreurs , R. Huiskes , J.
Bone Joint Surg 81 (1999) 1052–1057 .
[201] J.
Wu , R. Yan g , J. Zheng , L. Pan , X. Liu , Micro Nano Lett 14 (2019) 852–855 .
[202] S. Eosoly , D. Brabazon , S. Lohfeld , L. Looney , Acta Biomater 6 (2010)
2511–2517 .
[203] H. Doyle , S. Lohfeld , P. McHugh , Med. Eng. Phys. 37 (2015) 767–776 .
[204] H. Chung , H. Jee , S. Das , J. Mech. Sci. Technol. 24 (2010) 241–244 .
[205] M. Ziaee , A. Mahmood , N.B. Crane , J. Manuf. Mater. Proc. 4 (2020) 7 .
[206] Y. Du , H. Liu , Q. Yan g , S. Wang , J. Wa ng , J. Ma , I. Noh , A.G. Mikos , S. Zhang ,
Biomaterials 137 (2017) 37–48 .
[207] X. Huang , S. Lin , Y. Hu , Y. Liao , W. Wang , Z. Zhang , G. Rui , P. Cai , J. Biomater ,
Tissue Eng 8 (2018) 1038–1045 .
[208] L. Peña-Parás , J.A. Sánchez-Fernández , R. Vidaltamayo , in: Handbook of Eco-
materials, Springer, Cham, 2018, pp. 1–19 .
[209] J. Bai , R.D. Goodridge , R.J. Hague , M. Okamoto , Polym. Compos. 38 (2017)
2570–2576 .
[210] C.
Chua , K. Leong , K. Tan , F. Wiria , C. Cheah ,J. Mater. Sci. Mater. Med. 15
(2004) 1113– 112 1 .
[211] T. Kumaresan , R. Gandhinathan , M. Ramu , M. Ananthasubramanian , K.B. Prad-
heepa , J. Mech. Sci. Technol. 30 (2016) 5305–5312 .
[212] M. Ramu , M. Ananthasubramanian , T. Kumaresan , R. Gandhinathan , S. Jothi ,
Bio-Med. Mater. Eng. 29 (2018) 739–755 .
[213] B.P. Kamarajan , R. Murugan , A. Muthusamy , Int. J. Biomed. Res. 8 (2017)
467–474 .
215