ARTICLE IN PRESS
Biomaterials 27 (2006) 2651–2670
www.elsevier.com/locate/biomaterials
Review
Fabrication methods of porous metals for use
in orthopaedic applications
Garrett Ryan, Abhay Pandit, Dimitrios Panagiotis Apatsidis
National University of Ireland, Department of Mechanical and Biomedical Engineering, Nun’s Island, Galway, Ireland
Received 1 August 2005; accepted 7 December 2005
Abstract
Implant stability is not only a function of strength but also depends on the fixation established with surrounding tissues [Robertson
DM, Pierre L, Chahal R. Preliminary observations of bone ingrowth into porous materials. J Biomed Mater Res 1976;10:335–44]. In the
past, such stability was primarily achieved using screws and bone cements. However, more recently, improved fixation can be achieved by
bone tissue growing into and through a porous matrix of metal, bonding in this way the implant to the bone host. Another potentially
valuable property of porous materials is their low elastic modulus. Depending on the porosity, moduli can even be tailored to match the
modulus of bone closer than solid metals can, thus reducing the problems associated with stress shielding. Finally, extensive body fluid
transport through the porous scaffold matrix is possible, which can trigger bone ingrowth, if substantial pore interconnectivity is
established [Cameron HU, Macnab I, Pilliar RM. A porous metal system for joint replacement surgery. Int J Artif Organs 1978;1:104–9;
Head WC, Bauk DJ, Emerson Jr RH. Titanium as the material of choice for cementless femoral components in total hip arthroplasty.
Clin Orthop 1995;85–90].
Over the years, a variety of fabrication processes have been developed, resulting in porous implant substrates that can address
unresolved clinical problems. The advantages of metals exhibiting surface or bulk porosity have led researchers to conduct systematic
research aimed at clarifying the fundamental aspects of interactions between porous metals and hard tissue. This review summarises all
known methods for fabricating such porous metallic scaffolds.
r 2005 Elsevier Ltd. All rights reserved.
Keywords: Porosity; Bone ingrowth; Scaffold; Mechanical properties; Rapid prototyping
Contents
1.
2.
3.
4.
5.
Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .
Challenges for porous materials . . . . . . . . . . . . . . . . . . .
Closed-cell versus open-cell porosity . . . . . . . . . . . . . . . .
Fabrication techniques for closed-cell porous metals . . . . .
4.1. Random pore distribution. . . . . . . . . . . . . . . . . . .
4.2. Graded distribution of pores . . . . . . . . . . . . . . . . .
Fabrication techniques for open-cell porous metals . . . . . .
5.1. Non-homogeneous pore distribution . . . . . . . . . . .
5.1.1. Furnace sintered metal powders and fibres .
5.1.2. Space holder method . . . . . . . . . . . . . . . .
5.1.3. Replication . . . . . . . . . . . . . . . . . . . . . . .
5.1.4. Combustion synthesis . . . . . . . . . . . . . . . .
Corresponding author. Tel.: +353 91 492723; fax: +353 91 563991.
E-mail address: [email protected] (D.P. Apatsidis).
0142-9612/$ - see front matter r 2005 Elsevier Ltd. All rights reserved.
doi:10.1016/j.biomaterials.2005.12.002
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ARTICLE IN PRESS
G. Ryan et al. / Biomaterials 27 (2006) 2651–2670
2652
5.2.
Homogeneous pore distribution. . . . . . . . . . . . . . . . . .
5.2.1. Orderly oriented wire mesh . . . . . . . . . . . . . . .
5.2.2. Ferromagnetic fibre arrays . . . . . . . . . . . . . . .
5.2.3. Vapour deposition . . . . . . . . . . . . . . . . . . . . .
5.3. Functionally graded pore distribution . . . . . . . . . . . . .
5.3.1. Electrical field-assisted powder consolidation . .
5.3.2. Rapid prototyping . . . . . . . . . . . . . . . . . . . . .
6. Characteristics of porous metals that affect their performance .
7. Discussion and future work . . . . . . . . . . . . . . . . . . . . . . . . .
Acknowledgements . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .
References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .
1. Introduction
Despite the great progress that has been achieved in
orthopaedic biomaterials, fixation of implants to the bone
host remains a problem. Mismatch of Young’s moduli of
the biomaterials and the surrounding bone has been
identified as a major reason for implant loosening
following stress shielding of bone [1–3]. However, the
implanted material must be strong enough and durable to
withstand the physiological loads placed upon it over the
years. A suitable balance between strength and stiffness has
to be found to best match the behaviour of bone.
One consideration to achieve this has been the development of materials that exhibit substantial surface or total
bulk porosity in medical applications. The fabrication of
porous materials has been actively researched since 1943,
when B. Sosnik attempted to introduce pores into
aluminium by adding mercury to the melt [4]. In
biomedical applications the concept of using porous
materials has been investigated much later, whereby the
work of Weber and White from 1972 is one of the earlier
ones to mention the use of porous metals for osseointegration [5]. Numerous investigations into porous materials
where subsequently initiated in the early 1970s involving
porous ceramic [6,7], polymeric [8–12], and metallic
materials [13–16], which showed in animal studies to be
potential candidates for porous implants that would enable
bone ingrowth. Although ceramics portray excellent
corrosion resistance, the general opinion is that porous
ceramic structures, as they are available today, cannot be
employed as load bearing implants, due to their intrinsic
brittleness. Likewise, porous polymeric systems cannot
sustain the mechanical forces present in joint replacement
surgery. This led researchers to focus on porous metals,
based on orthopaedic metallic materials, as a consequence
of their superior fracture and fatigue resistance characteristics, which are required for load-bearing applications.
As aforementioned, a major problem concerning metallic implants in orthopaedic surgery is the mismatch of
Young’s modulus between bone (10–30 GPa) and bulk
metallic materials (between 110 GPa for Ti and 230 GPa
for Co–Cr alloys). Due to this mechanical mismatch, bone
is insufficiently loaded and becomes stress shielded, which
eventually leads to bone resorption. It has been suggested
that when bone loss is excessive, it can compromise the
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long-term clinical performance of the prosthesis [17]. It
may be responsible for implant migration, aseptic loosening, fractures around the prosthesis, and can pose technical
problems during revision surgery [17]. The relationship
between implant flexibility and the extent of bone loss has
been established in clinical patient series and animal
experiments and confirm that changes in bone morphology
are an effect of stress shielding and a subsequent adaptive
remodelling process [18–22].
Bobyn et al. performed bilateral non-cemented total hip
arthroplasties in canine models to determine the effect of
stem stiffness on stress-related bone resorption [18]. Two
partly porous femoral implants of substantially different
stiffness were designed for direct comparison. One was
manufactured from Co–Cr alloy, the other from titanium
alloy, but modified internally by drilling a hole that
extended from the stem tip to within 5 mm of the shoulder,
which transformed it into a hollow cylinder. Femora with
the flexible stems consistently showed much less bone
resorption than those with the stiff stems. Quantitative
analysis of paired cross-sections indicated an average of
25–35% more cortical bone area in femora that received
low stiffness hollow cylindrical stems. Similarly, Sychterz et
al. analysed the relation between femoral and implant
stiffness on bone remodelling based on cadaver specimens
from 20 patients with unilateral un-cemented hip replacements [22]. They showed that axial bone stiffness was the
variable most strongly correlated to bone loss and that the
stem to bone stiffness ratio accounted for 46% of the
variance in bone loss data.
Titanium and its alloy (Ti6Al4V) have elastic moduli less
than 50% of that in Co–Cr implants so that their use would
help reduce the extent of stress shielding. Although
fabrication of implants from materials with lower elastic
moduli can reduce stress shielding the stiffness mismatch to
bone is still substantial [23]. A suggestion to overcome this
drawback could be the use of porous materials in stems.
The clinical literature of the past 30 years records a variety
of approaches to this end and several researchers have
performed studies aimed at clarifying the fundamental
aspects of interactions between porous metals and hard
tissue. Porous materials in arthroplasty implants are
increasingly attracting the widespread interest of researchers as a method of reducing stiffness mismatches and
achieving stable long-term fixation by means of full bone
ARTICLE IN PRESS
G. Ryan et al. / Biomaterials 27 (2006) 2651–2670
ingrowth and there have been a number of previous reviews
on the many different porous coatings and fully porous
matrices that have been developed [11,24–33]. The present
review article aims at summarising the most common
methods for fabricating porous materials and describe
them in light of their suitability for use in orthopaedic
implants. New research into local and systemic factors that
enhance bone ingrowth fixation is also reviewed, as are
initial problems that have so far been associated with
implants based on porous materials.
2. Challenges for porous materials
A major concern with the use of porous implants in
highly loaded applications is the effect the porous matrix
might have on fatigue strength. Studies have shown that
both Co–Cr alloys and Ti–6Al–4V alloys experience drastic
reductions in fatigue strengths when fabricated as porous
coatings on solid core structures [34–37]. It has been shown
that the high cycle fatigue strength of porous coated
Ti–6A1–4V alloy is approximately one-third that of the
solid alloy equivalent shape, probably even less in fully
porous matrices [38]. A study by Yue et al. revealed crack
initiation in the substrate close to the particle-to-solid core
sinter neck region, using scanning electron microscopy, and
it is concluded that stress intensification due to these
interface regions are major sources of weakness with
respect to fatigue strength. The bond sites between the
coatings and implant have irregular geometries that can act
as stress concentrations. This is sometimes referred to as
the notch effect. This notch effect is a localised condition
that affects implant strength in the region of the porous
coating [34]. To achieve a functionally strong implant,
porous implant design needs to account for these losses in
material strength.
The current practice in designing porous titanium alloy
implants is to avoid porous coatings on surfaces that will
be subjected to significant tensile stresses in vivo [39]. Cook
et al. showed that an approximately 15% improvement in
fatigue properties of porous Ti–6Al–4V could be achieved
through post-sintering heat treatments that produce
microstructures that are more resistant to crack initiation
and propagation [40]. Also by modelling porous-coated
implants using linear elastic, plane strain finite element
analysis Wolfarth et al. predicted a doubling of fatigue
strength when optimising conventional porous geometries
[38]. Mechanical properties of porous materials can be
altered and optimised by controlling porosity, pore size
and shape as well as pore distribution. This review outlines
the various fabrication methods of metal foams and
provides an indication as to how far they are able to
provide control over these parameters.
It is commonly accepted that, in the long term, total joint
replacement is associated with adverse local and remote
tissue responses that are mediated by degradation products
of prosthetic materials [41]. Particular interest has centred
on the metal-degradation products of total joint replace-
2653
ments because of the known toxicity of the various
elements in those alloys used for implants. Corrosion can
also severely limit the fatigue life and ultimate strength of
the material, leading to the mechanical failure of the
implant. There is a low but finite prevalence of corrosionrelated fracture of implants [42]. Increased surface areas,
such as in porous implants, have shown higher corrosion
rates when tested in vitro compared to conventional non
porous-coated implants [43,44]. This has caused concerns
regarding long-term safety of porous implants. Enhanced
metal ion release could increase the probability of metal
sensitisation and associated allergic responses in individuals could increase the susceptibility to tumour formation
[45]. This matter would have to be addressed by only
implanting surface-treated porous materials into the body.
Available techniques to modify implant surfaces are
reviewed at the end of this review article.
Becker et al. studied the corrosion behaviour and
mechanical properties of three medical grade alloys; 316L
stainless steel, the Co–29Cr–6Mo and Ti–6Al–4V alloys,
manufactured using the cold compaction and sintering
route to contain two grades of porosity (30% and 10%)
[46]. Ti–6Al–4V performed best in the combination and
was therefore the recommended material for use in porous
implants. A study by Seah et al. investigated the influence
of pore morphology on corrosion [47]. They investigated
the corrosion resistance of porous titanium parts that were
fabricated with powder metallurgy methods and had
varying porosities and pore sizes. Again, 316L exhibited
poor corrosion resistance compared to titanium. Their
main finding was that corrosion resistance decreased with
decreasing porosity, which was attributed to the small,
isolated pore morphology that traps ionic species and
restricts the access of oxygen, which in turn limits the
available oxygen for the creation formation of important
corrosion resistant passive layers. In highly porous
compacts with an open, interconnected pore morphology,
the free flow of species resulted in a material with increased
corrosion resistance.
Static stresses were also found to affect corrosion
behaviour of porous materials. In a study by Bundy et al.
polished, grit-blasted, and porous-coated surfaces were
investigated for their susceptibility to stress-enhanced ion
release (SEIR) [48]. The porous-coated materials were
shown to be the most susceptible to SEIR, which was
induced by ex vivo elastic loading of the materials. Stress
affected a number of corrosion parameters, but lowering of
breakdown potentials and increase of corrosion currents
were the most relevant to the clinical situation. It can be
assumed that similar effects will be true for fully porous
materials, where the total available surface area is even
further increased. Assuming that similar effects occur in
vivo under load application, then tests on unstressed alloys
in vitro could be grossly underestimating the potential in
vivo ion release rates.
From the above it becomes apparent that there are
numerous known- and possibly several unknown-factors
ARTICLE IN PRESS
G. Ryan et al. / Biomaterials 27 (2006) 2651–2670
2654
Table 1
Summary of various fabrication methods for porous metals and their categorisation according to the resulting pore distribution
Fabrication methods
Closed-cell
Open-cell
Random pore distribution
Graded pore distribution
Non-homogeneous
Homogeneous
Functionally graded
Gas injection into the
metal melt [52]
Plasma spraying [15]
Sintered metal powders
[62]
Sintered metal fibres [69]
Orderly oriented wire
mesh [83]
Vapour deposition [92]
Rapid prototyping [55]
Space holder method [70]
Ferromagnetic fibre array
[85]
Rapid prototyping [55]
Decomposition of
foaming agents [53]
Replication [74]
Combustion synthesis [81]
Plasma spraying [15]
that need to be considered in the fabrication of porous
materials and the methods that are available for achieving
these require continuous improvement, in order to
adequately address these needs. This article summarises
the existing techniques and puts them under the light of
their potential performance in vivo.
3. Closed-cell versus open-cell porosity
A major classification of porous metals, or metal foams,
is between open-cell and closed-cell. In closed-cell foams
each cell is completely enclosed by a thin wall or membrane
of metal, whilst in open-cell foams the individual cells are
interconnected, allowing tissue to infiltrate the foam and
anchor it into position. Closed-cell porous metals are
usually the result of a random foaming process, in which
the size, shape and location of pores within the matrix
varies, depending on the parameters of the fabrication
process. The result is usually a porous material with limited
porosity and, often significant, variations in pore size and
shape, although careful selection of the foaming parameters can improve homogeneity [49].
It is recognised that there are three distinct types of
porous implants: (1) partly or fully porous-coated solid
substrates; (2) fully porous materials; (3) porous metal
segment joined to a solid metallic part. There are several
applications that can potentially use both porous-coated
and fully porous implants. These include:
spinal fixation devices,
fracture plates,
wires, pins and screws,
artificial ligament attachment implants
cranio-facial implants,
maxillofacial implants,
bone graft material to fill tumour defects.
Implants with solid cores and porous coating structures
are more appropriate when the porous metal alone does
Electro discharge
compaction [103–105]
not provide sufficient mechanical strength to sustain the
physiological loads, such as in:
dental implants,
joint arthroplasty implants.
Moreover, the fabrication of open-cell porous metal
implants can be divided into three categories, classified
according to the state the metal is processed in:
1. solid state in powdered or fibre form,
2. liquid metal,
3. metal vapour or gaseous metallic compounds.
Different processes vary in complexity of preparation and
also in the type of porous material that they produce. Thus,
some processes such as casting or vapour deposition
techniques tend to allow greater control over pore size,
distribution and interconnectivity. Other processes involving
decomposition of foaming agents in either molten or powder
metal matrices give lower porosities and less predictable pore
distribution and interconnectivity. The former can produce
open-cell geometries, whereas the latter usually result in
closed-cell matrices. In the following section, we outline
some of the main methods for fabricating fully porous
metals. We demonstrate how the processes have evolved
from basic sintering of metal powders or fibre particles to
complicated procedures that can yield foams with higher
porosity, interconnectivity, more uniform pore size and
greater control over pore distributions. The reviewed
methods are summarised in Table 1 and categorised into
closed-cell and open-cell structures of random or orderly
pore distributions. Some of the mentioned techniques could
potentially produce more than one category of porous
material and are therefore listed under different headings.
4. Fabrication techniques for closed-cell porous metals
Most industrial engineering applications that require the
use of porous metals have come to depend heavily on
ARTICLE IN PRESS
G. Ryan et al. / Biomaterials 27 (2006) 2651–2670
closed-cell porous metals with optimised structures and as
few impurities as possible, in order to provide adequate
mechanical properties. However, when it comes to
fabricating functional porous metals it is open-cell porous
metals that are the preferred standard [50]. Such a
functional role for the use in implants is of course the
facilitation of bone ingrowth, for which open-cell materials
are required. This property in combination with the
low stiffness of open-cell porous metals is the reason for
maintaining the focus in this review on this category
of porous materials. Nevertheless, closed-cell porous
metals can still play a role in orthopaedic implants, as is
discussed in the following section, and are thus also
reviewed.
4.1. Random pore distribution
Although closed-cell metal foams do not allow for much
bone ingrowth, due to the large numbers of isolated pores,
they still possess the potential to be used in orthopaedic
load bearing implants. In particular, the lowered structural
stiffness brought about by the introduction of voids allows
the reduction of their bulk stiffness, enabling a match with
the mechanical properties of bone, which in turn reduces
stress shielding of the bone host. Thereby, the pore size,
shape and distribution are of importance only because of
the mechanical strength and fatigue resistance of the
porous metal that is associated with these parameters.
Thereby it has been stipulated that non-homogeneous
distributions of pores as well as localised thin cell walls can
lead to a reduced materials strength and early fatigue
failure [51]. Fixation of implants based on closed-cell
porous metal would have to be achieved either by using
PMMA bone cement or allowing bone ingrowth onto an
additionally fabricated porous coating.
At present, there are a number of methods used to
fabricate closed-cell foams. There are two general routes to
generate porosity: melting and powder metallurgy. In the
first, self foaming structures are manufactured either by gas
injection through the melt (Cymats/Hydros), or by the
addition of gas forming elements into the liquid metal
(ALPORASs) [52]. These methods have been used to
fabricate Al, Zn and Mg foams, however, they are
unsuitable for the manufacture of Ti foams, due to the
high melting temperature and the associated reactivity of
Ti with oxygen residues in the ovens. In the powder
metallurgy approach the structures are obtained either by
sintering hollow spheres or by melting of powder compacts
that contain a gas evolving element such as TiH2 (Alulights/Foaminals) [53]. This approach has been known to
yield a relatively homogenous structure and can be used in
the manufacture of high melt metals and alloys. Fatigue
strength can be improved by incorporating an adequate
mixing strategy of the metal and foaming agent powders
due to a resulting homogenous pore distribution. This
helps minimising stress concentrations within the structure
and increasing fatigue life significantly.
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4.2. Graded distribution of pores
A common technique for producing porous structures is
plasma spraying. It can be used to create rough solid
surface textures, porous surface coatings on solid cores and
also fully porous structures [54]. The former of these three
types of structures that can be produced by the plasma
spraying method are often regarded as porous, although in
the context of bone ingrowth they probably merely
improve surface anchoring of bone tissue compared to
highly polished surfaces. In the contrary, porous surface
coatings on solid cores and fully porous structures whether
they are open-cell or closed-cell, allow bone ingrowth into
pores. A schematic description of the plasma spraying
process is shown in Fig. 1.
During plasma spraying, an electric arc is generated
between two water-cooled electrodes in a gun. The arc
heats a gas to extremely high temperatures (up to
20,000 1C), partially ionising it and forming a plasma jet.
The gases are accelerated by the tremendous expansion in
volume and pass through the jet-shaped anode at a high
speed. The powder for the coating is injected into the
plasma gas stream, using a carrier gas where they are
accelerated to a high speed, melted and impacted onto the
substrate with high kinetic energy. Porous coatings with
varying degrees of porosity can be created on the substrate
by adjusting the spraying parameters (Fig. 2). Plasma
spraying is normally performed under vacuum where
interactions between the plasma jet, powder, substrate
and the surrounding atmosphere are reduced significantly.
However, another variant of the process is reactive plasma
spraying, where the starting powder materials are reacted
with inert or reducing gaseous species and introduced into
the plasma jet. Due to the short reaction time, the
deposited material is usually a composite of the starting
material forming the matrix and in situ synthesized phases.
This is indispensable for plasma spraying of titanium,
which is a sensitive material prone to oxidation and
nitrogen absorption [54].
Fig. 1. Schematic representation of the plasma spraying process.
Changing the spraying parameters varies the porosity of the resulting
coating layer.
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of porous coating and solid core, whereby materials with
thicker porous coatings result in a lower composite
stiffness than those with smaller porous coating thickness.
However, the techniques mentioned in these references
involve at least a two-step-processing route that increases
the cost of the product. Other problems include contamination and the presence of impurity phases.
5. Fabrication techniques for open-cell porous metals
5.1. Non-homogeneous pore distribution
Fig. 2. Cross section of the interface between substrate and coating as
produced by plasma spraying. Reprinted with permission from [54].
Hahn and Palich first described titanium plasma-sprayed
coating for fabricating porous-coated implants [15]. They
used titanium hydride powders fed into a plasma flame,
whereby the decomposed titanium was deposited onto an
appropriate substrate (titanium hydride starts to decompose at around 600 1C and reaches complete dissolution
into titanium and hydrogen at 1000 1C). The report of their
early studies described the use of a carrier gas consisting of
hydrogen (5–15%) and nitrogen (balance) for the plasma
spraying. By choosing an appropriate gun-to-substrate
distance, a thin coating (approximately 900 mm thick) with
porosity that varied from zero at the substrate interface to
about 50% at the coating surface was formed. The
formation of other metal surface coatings such as Co–Cr
alloy, stainless steel or Ti–6Al–4V by this process is also
possible. However, coatings prepared with this method
result in irregular porosities and the pore interconnectivity
is quite low compared to other techniques.
Nevertheless, graded porous titanium coatings have been
produced using plasma spraying and are characterised by a
gradual change in porosity from the substrate-coating
interface to the coating surface [55,56]. The graded porous
coatings consisted of three layers. The outer layer was full
of macro pores with a surface roughness of approximately
100 mm. The diameter of many macro pores reached and
even surpassed 150 mm, which is beneficial for bone to grow
into the coating. The middle layer consisted of a mixture of
micro pores and macro pores. The inner layer was a very
dense and tight interface layer that included mechanical,
physical, and metallurgical bonding. This gradual change
in porosity means that the Young’s modulus of the
material is better adapted to the elastic properties of bone
compared to solid metals, thus reducing stress shielding
between the implant and the bone. This is largely due to the
reduced stiffness of the porous coating, when compared to
the solid core’s stiffness. It is evident, of course that the
composite material stiffness depends on the thickness ratio
5.1.1. Furnace sintered metal powders and fibres
The simplest fabrication technique for making metallic
foam is based on the partial densification during sintering
of metal powders. This technique is known as powder
metallurgy and is a mature metal-forming technology used
in the fabrication of both porous coated and fully porous
metallic implants. It is the technology of producing
materials by compacting, binding, and sintering metal
powders. The sintering operation is in essence a hightemperature treatment process that causes the powder
particles to bond to each other with only minor change to
the particle shape. A microscopy image of a sintered
powder sample is shown in Fig. 3. A binder is often added
with the particular purpose of holding powder particles
together. This ensures a greater area for mass transport
between the particles in the solid-state diffusion process.
Fig. 3. Cross-section of an open-cell metal foam that was produced by
partial sintering of commercially pure Ti powders, showing sharply cusped
pores (porosity is 24%). Reprinted with permission from [58].
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The volume fraction of porosity is associated with the
degree of particle interconnectivity and particle size. It can
be controlled by process variables such as compacted
powder density, sintering temperature and time, and
alloying additions. The limitation of the powder sintering
approach is that pore size and shape are dictated by the
powder size and shape. For spherical powder particles the
porosity is limited to 50% and the shape of the pores is
highly non-spherical [57].
Powders based on Co–Cr alloys [59,60], commercially
pure Ti [61,62], and Ti alloys (mainly Ti–6Al–4V [63] and
NiTi [64]) have been used to create both porous surfaced
and fully porous implants, using particles ranging in size
from 50 mm to 1 mm. Oh et al. sintered spherical unalloyed
titanium powders with and without applied pressure and
achieved a porosity range of 5–37% [62]. Young’s modulus
and compressive yield strength decreased linearly with
increasing porosity, and at 30% porosity, the stiffness of
the porous titanium was close to that of human cortical
bone (20 GPa). Thieme et al. produced titanium foams with
a gradient in porosity by sintering a stack of three powder
layers with different particle sizes and silicon contents [55].
Silicon is added to produce a transient liquid phase during
sintering, resulting in a substantial increase of the particle
neck geometry whilst at the same time reducing porosity.
The volume fraction and pore size varied from 22% and
48 mm for the finer powder layer to 45% and 200 mm for the
coarser powder layer. The Young’s modulus of uniform
non-graded stacks ranged from 5 to 80 GPa, as determined
by ultrasound velocity measurements.
The sintering operation involved in making titanium and
titanium alloy implants requires a non-oxidising environment to achieve good bonding, which typically means the
need for a high vacuum oven (105 mbar) and sintering
temperatures of around 1250 1C. Particle contamination
(by oxidation or some other surface contaminant) would
hinder particle bonding. The pore size, volume fraction,
morphology and distribution throughout the sample
thickness and the inter particle neck size have a major
impact on the mechanical properties of the resulting
material. Sintered metal powders are often very brittle
with poor toughness and are prone to crack propagation at
low stresses or at low impact energies. Under fatigue
conditions, cracks are likely to initiate at the sintered necks
of individual powder particles. With sintered coatings the
aim is to provide strong bonds between each of the powder
spheres (beads) and between the coating and the implant
without significantly degrading the strength and corrosion
resistance of the component.
Sintered fibre metal coatings can be another alternative
for cementless implant fixation in load bearing applications. The process used to create these coatings is akin to
that described for sintered powder structures. The points of
contact between fibres become metallurgical bonds during
the sintering process and the aggregate acquires considerable mechanical strength. The pores are fully interconnected so that growing bone can permeate the entire
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composite. Although stainless steel and Ti fibre structures
have been described in literature [43,65–68], only the Ti
system is used clinically [69]. A major disadvantage of the
metal fibre sintered porous coatings compared to the
coatings made by powder metallurgy techniques is that
fibres must be compacted to a form prior to sintering. This
means it is difficult to coat complex shapes that do not
allow compaction forces to be applied directly onto fibres
overlying the substrate surface. The problem of ensuring
good bonding between the coating and the substrate lies in
that the fibres may spring back during metal fibre
compaction, and can result in inadequate coating-substrate
contact leading to regions of poor bonding after sintering.
Although some literature describes sintered fibre metal
specimens as mechanically more stable than powder metal
structures [65], it can be argued that it is the mode of failure
that differentiates these structures from others. This means
that the fibres do not fail mechanically by the propagation
of a crack, but only over a large displacement range by
means of tearing, which represents the progressive rupture
of the bonds between fibres or of individual fibres.
However, appropriate sintering procedures can ensure
integrity of both powder-made and metal fibre sintered
coatings. Unfortunately, the porosity of metal fibre coatings is limited to 30–50% by volume, which directly limits
the maximum interfacial strength that can develop by bone
ingrowth.
5.1.2. Space holder method
The space holder method is a fabrication process that
can produce porous metal samples of greater porosity. A
schematic showing the relevant processing steps is shown in
Fig. 4. The process begins by mixing the metal powders
with an appropriate space holder material and is followed
by the compaction (e.g. uniaxial or isostatic) of the mix to
form a green body. The resulting pellet is then subjected to
a low-temperature heat treatment process that is designed
to remove the space holder, which also leads to initial stage
sintering of the metal particles that are in contact. The
result is an initial neck formation. Continued sintering at
higher temperatures further develops sinter neck growth.
This leads to densification of the structure and associated
improvement of structural integrity.
In the preparation process, the powder size of the metal
powder should be smaller than the average powder size of
the space holder. In addition, the compaction pressure for
the metal powder and space holder mix must be high
enough to give the structure sufficient mechanical strength
so that it will retain its geometry throughout the foaming
process. This method provides a foamed structure with a
close to homogenous pore structure and high levels of
porosity (60–80%) [70]. By choosing the size, shape and
quantity of the space holder used the mechanical properties
of the metal foam can be adjusted. Smaller sizes of the
space holder particles can be obtained by sieving. An
example of the porosity that can be achieved using this
process is shown in Fig. 5.
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A general difficulty of this method is the removal of large
quantities of the space holder materials from the compacted mix. Bram et al. used carbamide (urea) powders,
which could be removed at temperatures below 200 1C,
with minimal contamination of the titanium powders [70].
Subsequent sintering at 1400 1C for 1 h resulted in foams
with porosities in the range of 60–70% and pore diameters
in the range of 0.1–0.24 mm, depending on the carbamide
particle size. The compressive and bending mechanical
properties for these foams, and a plateau stress of
approximately 10 and 100 MPa for porosities of 77% and
60%, respectively, was reported. Wen et al. used ammonium hydrogen carbonate as space holder, which decom-
Fig. 4. A sequence of individual processing stages are required in the
powder metallurgical space holder method.
Fig. 5. SEM micrograph of the porous titanium foam fabricated using the
space holder method. Reprinted with permission from [71].
poses at 200 1C and a sintering temperature of 1200 1C for
2 h. The resulting foam had a porosity of 78% and
exhibited a compressive strength of 35 MPa, whilst
Young’s modulus was 5.3 GPa. These values were a close
match to those for cancellous bone, 2–10 MPa and
1–10 GPa, respectively [72].
A similar approach was taken by Tuchinskiy et al.,
whereby rods consisting of a shell and core were compacted
using a specially designed die [73]. The shell consisted of a
mixture of titanium powder and polymer binder and the
core consisted of organic filler, which could later be
thermally removed. The compacted rod was then cut into
predetermined lengths that were subsequently poured into
a die and compressed. Fillers and binders were then
removed from the resulting green pellet prior to sintering.
This resulted in the generation of cylindrical pores in the
foam. The ratio of the core area to the shell area in the rod,
the compression conditions and the sintering conditions all
affected the final porosity of the foam. A major aspect of
the die’s design, however, was that the material could be
interchangeably extruded instead of being compacted by
the die, which would mean that the pores would be
oriented predominantly in one direction. The result of this
was that by varying the orientation of the pores and the
length of the cut rods, it was possible to create anisotropic
porous structures.
5.1.3. Replication
An approach that is related to the above technique uses a
three-step procedure, as depicted in Fig. 6, for the
production of highly porous materials. Li et al. utilised
this method to produce porous titanium and titanium alloy
structures [74]. Polyurethane foams were immersed in
titanium slurry comprising Ti–6Al–4V powder (70%wt),
H2O (20%wt) and ammonia solution. The ammonia
solution was added to improve the rheological properties
of the slurry. The sample was subsequently dried and the
process was repeated until all the struts of the polyurethane
foam were coated with Ti–6Al–4V powder. After thermal
removal of the polyurethane scaffold and binder and
subsequent sintering of the powders, a reticulated open-cell
foam with hollow titanium struts remained (Fig. 7).
Three types of pores are present in these foams, which
are in order of increasing size: primary porosity on the
surface of the hollow struts, secondary porosity at the core
of the hollow struts previously occupied by the polyurethane foam, and open tertiary porosity between struts.
Fig. 6. Schematic representation of the three-step replication process. A pattern of the final design of the porous metal is made from a different material
and reproduced with the actual desired material via an intermediate infiltration step.
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The porous Ti-alloy made in this way was found to possess
a porosity of 88% and compressive strength of 10 MPa. It
was found that the rheological properties of the Ti slurry
play an important role in the impregnation process, which
is governed by the particle size and shape of the raw
powder, the type and content of the binder, the solid/liquid
ratio, pH, sedimentation behaviour of the slurry and the
amount of air bubbles in the mix. It was also discovered
that a rapid drying process is important in maintaining a
positive replica shape. It was later shown that a second
deposition of powder slurry on a previously sintered foam
followed by a second sintering resulted in increased density
and improved compressive strength (80% and 36 Mpa,
respectively), as a result of the correction of flaws in the
titanium struts created during the first cycle [75].
5.1.4. Combustion synthesis
A recently developed effective method of producing high
purity porous alloys and in particular nickel titanium
alloys is a process known as combustion synthesis (CS). In
CS particle fusion is obtained through an extremely rapid
self-sustaining exothermic reaction driven by the large heat
released in the synthesis. The reactants, in the form of fine
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powders are usually dry-mixed and cold pressed. The
exothermic reaction can then be instigated under two
different regimes: (a) thermal explosion mode, in which the
reactants are gradually heated until reactions take place
simultaneously throughout the whole sample, and (b) selfpropagating high thermal synthesis (SHS), which is
characterised by the fact that once ignited, a strong
exothermic reaction propagates as a combustion wave
through the entire mixture, without requiring additional
energy. The samples are placed in a controlled and inert
atmosphere and ignited by means of an electrically heated
coil, a laser beam, or an electric discharge. Various
processing parameters such as the reactant particle size of
powder, the use of a binder and the compaction pressure
affect the final microstructure and porosity of the sample
[76]. The mechanical properties of the resulting structure
are controlled by means of taking these parameters into
account. A schematic of the stages in this process is shown
in Fig. 8.
Several researchers have investigated the use of CS in the
fabrication of porous nitinol (NiTi) [76–80]. Li et al.
successfully prepared porous NiTi shape memory alloys
with anisotropic pore structure by SHS [81]. Titanium
(21 mm) and nickel (38 mm) powders were mixed and cold
pressed and then placed in a reaction chamber where they
were ignited at a preheating temperature of 550 1C using a
tungsten coil heating element, based on the following
reaction:
Ni þ Ti ! NiTi þ 67 kJ mol1 :
Fig. 7. SEM micrograph of reticulated Ti–6Al–4V foam produced by
sintering of powders deposited onto a temporary polyurethane scaffold.
Reprinted with permission from [74].
Due to the relatively low exothermic characteristic of the
reaction between Ni and Ti, preheating the sample prior to
the ignition was necessary to achieve self-sustained
combustion. The pores in the compacted pellet (40% vol),
the transient liquid phase, the volatilisation of impurity,
and the escape of the adsorbed gases all contributed to the
final porosity of the compact, which was finally 54% vol. A
structure of channels in the compact formed along the
propagation direction of the combustion wave. This can be
seen in Fig. 9, which shows a porous sample made by this
process.
Fig. 8. Schematic representation of the stages that are involved in combustion synthesis. The compaction acts as the trigger for initiating an explosion,
which synthesises the mix as it propagates.
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Fig. 9. Scanning electron microscope image of porous titanium–nickel
material fabricated by self-propagating high thermal synthesis. Reprinted
with permission from [82].
An advantage of this fabrication process is the high
purity of the resulting foams, which is largely due to the
expulsion of volatile impurities under the extremely high
temperatures in the process. In vivo studies suggest
strongly that there is natural bone ingrowth with no
apparent immune response in the composite materials
produced in this process [76].
5.2. Homogeneous pore distribution
5.2.1. Orderly oriented wire mesh
Orderly oriented wire mesh (OOWM) coatings for
orthopaedic implants were created to provide for a large,
uniform pore size for tissue ingrowth [83]. Like the fibre
mesh structures, this system makes use of small-diameter
metal wires. However, rather than cutting these to short
lengths for compaction, the continuous wire lengths are
woven into a regular meshwork that is subsequently
pressure sintered onto the solid substrate. The wire
diameter, inter-wire spacing, and geometric distribution
of the wires determine the dimensions of the interconnecting porosity. Prior to sintering, the woven wire mesh is precompacted in order to increase the contact zones between
the wire mesh and substrate. As with the fibre mesh
systems, the OOWM coatings are more conveniently
applied onto flat implant surfaces. For larger implants
such as hip and knee joint replacements, this system should
be adequate provided the wire meshes are placed in
position where bone contact ingrowth can be ensured.
The OOWM system was designed to avoid the difficulties
of particle and fibre detachment that could occur with the
powder metallurgy method for fabricating porous-coated
implants and metal fibre composite implants. In addition
to the regular weaves, twill weaves can be used for porous
coatings (Fig. 10).
Twill weaves are made by passing the wires consecutively
above and below two crossing wires. Furthermore, two
parallel neighbouring wires pass from above to below at
Fig. 10. Micrographs of twill weave triple OOWM coating (a) crosssection ( 50), and (b) scanning electron view ( 20). Reprinted with
permission from [83].
one crossing wire apart, that is their weave is shifted by one
wire. The desirable aspects associated with the twill weave
are ease of fabrication of the mesh and greater flexibility
and formability in subsequent coating procedures.
5.2.2. Ferromagnetic fibre arrays
Markaki et al. suggested that bone growth could be
stimulated in vivo via a magneto-mechanical mechanism
that involves the transmission of stresses and strains to
growing bone via small local deflections within a porous
ferromagnetic material, induced by the application of an
external magnetic field [84]. It has been determined that
strain levels of at least about 1 millistrain (0.1%) are
needed, in order to stimulate bone growth. Utilising this
fact, they have created implants with an outer layer of
highly porous ferromagnetic fibres bonded together [84].
The porous specimens were created by spraying a small
quantity of fibres with a slow setting aerosol glue and then
sprinkling some braze powder over them. The fibres, with
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stiffness of this structure, a suitable prosthesis would
require a dense metallic core with the porous fibre array
bonded to the outer layer, giving an overall stiffness of the
component equal to that of bone.
Fig. 11. Scanning electron micrograph of a bonded fibre array material,
made by brazing ferromagnetic stainless steel fibres. Reprinted with
permission from [85].
braze particles adhering to them, were packed into a long
quartz tube and brazing was carried out at 1200 1C.
Resulting specimens had porosity levels of about 75–90%
and the fibre array had average pore sizes of 100–300 mm.
A typical fibre content and distribution is illustrated in the
scanning electron micrograph shown in Fig. 11.
The primary requirement in the above technique is that
the fibre material is ferromagnetic. It has been established
that several such materials exhibit good corrosion resistance in biological fluids [86–89]. In fact, it has been shown
that ferromagnetic stainless steels display superior resistance to pitting and crevice corrosion, and to the leaching
of iron, chromium and nickel, compared to 316L, a
common non-magnetic stainless steel in use for implants
[89].
Markaki et al. developed a simple analytical model to
predict the expected levels of deformation [84]. When a
magnetic field is applied, the fibrous array deforms
elastically, as a result of the tendency for such fibres to
align with the field. In-growing bone tissue that is gradually
filling the inter-fibre space would be mechanically strained
during this deformation. Predictions indicate that ingrowing bone tissue, with a stiffness of around 0.01–
0.1 GPa, could be strained to beneficial levels (1 millistrain), using magnetic field strengths similar to those used
in current diagnostic procedures (1 T). Experimental
measurements confirm the broad validity of the model,
although no work has been done so far involving bone
tissue. Due to the high fibre segment aspect ratio needed to
confer the potential for magneto-mechanical induction of
therapeutically beneficial strains and the resulting low
5.2.3. Vapour deposition
Recently, new metallurgical techniques have been
attempted for creating metallic matrices of much greater
porosity. One example of these is Trabecular Metals
(Hedrocels, Implex corp.), which has a pore volume
fraction of 75–85% and is characterized by continuous
interconnecting pores, or cells, each of which possess the
shape of a dodecahedron. The fabrication of Trabecular
Metals (Hedrocels, Implex corp.) is based on a medical
grade polyurethane foam.
The cellular density and pore geometry of the polyurethane foam are the precursors to the pore size and
homogeneous mechanical properties of Trabecular Metals. The polyurethane foam is reticulated and subsequently pyrolysed, resulting in a low-density reticulated
vitreous carbon (RVC) skeleton that can be machined or
crushed into a variety of shapes or pre-forms. The vitreous
carbon pre-forms are in turn transformed into Trabecular
Metals by means of a proprietary chemical vapour
deposition (CVD), during which commercially pure tantalum is deposited throughout the RVC pre-forms [90].
CVD is the generic name for a group of processes that
involve depositing a solid material on a substrate by
activating the reactants in the gaseous phase where they
react chemically. Fig. 12 illustrates some of the fundamental aspects of this process. Reactant gases, often
diluted in a carrier gas, at room temperature enter the
reaction chamber and the gas mixture is heated by
radiation as it approaches the deposition surface [90] or
placed upon a heated substrate. Depending on the process
and operating conditions, the reactant gases may undergo
homogeneous chemical reactions in the vapour phase
before striking the surface. Near the surface, thermal,
momentum, and chemical concentration boundary layers
Fig. 12. Schematic of the CVD process, which involves the use of an
intermediate reticulated vitreous carbon substrate.
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Fig. 13. Scanning electron micrographs of porous tantalum, showing the cellular structure formed by the tantalum struts. Reprinted with permission from
[91].
form as the gas stream heats then slows down due to
viscous drag, eventually changing the chemical composition. Heterogeneous reactions of the source gases or
reactive intermediate species (formed from homogeneous
pyrolysis) occur at the substrate surface forming the
coating layer. Gaseous reaction by-products are transported out of the reaction chamber.
The relative density, or conversely the porosity, is
controlled by the duration of the CVD process. This
dictates the structural properties of the final product. The
porosity (pore volume) of a typical Trabecular Metals
component is approximately 75–85%, and is characterized
by an average pore size (diameter) of between 550 mm [92].
The porous material is comprised of approximately 99%
tantalum and 1% vitreous carbon, by weight. As shown in
Fig. 13 the appearance of Trabecular Metals approximates that of cancellous bone with regard to the open,
interconnected porosity and strut thickness inherent to
cancellous bone. Each pore is defined by the tantalum
struts, which intersect at nodes that typically include four
struts per junction.
The strength and stiffness of porous materials increases
with decreasing porosity, which is typical of porous solid
materials such as cancellous bone [93]. The flexural rigidity
of porous tantalum was recently found to increase with
relative density to the power of 1.2, and was found to be
similar to that of the human fibula [94]. The material is
found to have many superior mechanical properties to
other materials used as a hard tissue scaffold [92,95,96].
Other tests have shown that variability of the foam
structure and carbon strut dimensions coupled with
variability in the layered structure and thickness of the
tantalum, probably account for the significant deviations
noted during mechanical characterisation [97], however,
the ductility of this material was found to be far superior to
other porous materials such as bone, other naturally
occurring materials, ceramics and composites with ceramic
fillers.
Bone ingrowth characteristics of porous tantalum
vertebral body fusion devices were investigated by Zou
et al. in porcine models and compared to carbon fibre
vertebral body devices [98]. This study demonstrated that
Fig. 14. Schematic of FAST, comprising a die and punch fixture and an
external current supply for the heating and consolidation of the powder
mix.
the radiographic and histological appearance of the porous
tantalum was equivalent to the carbon fibre cage, which
indicates that the difference in osteoconductivity was not
due to the tantalum and carbon material.
5.3. Functionally graded pore distribution
5.3.1. Electrical field-assisted powder consolidation
The general term of electrical field activated sintering,
involves the application of an external current to assist
powder consolidation. In recent times, there has been a
growing interest in this application to create porous
implants for orthopaedic applications. The technique is
known under different names, such as field assisted
consolidation technique (FAST) [99], spark plasma sintering (SPS) [100,101], plasma activated sintering (PAS) [102],
and electrical discharge compaction (EDC) [103–105]. All
techniques have in common the combination of an
electrical discharge with rapid heating and pressure
application to achieve fast sintering of powders.
A schematic showing the FAST process is shown in
Fig. 14. The equipment consists of a uniaxial compression
device and the electrical components to apply the pulsed
and steady currents. The powder of choice is directly
loaded into a punch and die unit. Graphite die and graphite
punches are commonly used. The graphite confinement provides a reducing component to the sintering
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G. Ryan et al. / Biomaterials 27 (2006) 2651–2670
environment. The device is commonly placed in a vacuum
or gas controlled environment.
The consolidation process consists of two stages: (1) an
initial activation through the application of a pulsed
voltage, and (2) the subsequent heating and densification
by using DC current. A typical pulse discharge is achieved
by the application of a low voltage (10 V) and a
600–1000 A current. The basic difference between FAST
and EDC is in the number of discharges: in EDC the
electrical energy is suddenly released by discharging a
capacitor bank through the powder compact whereas
multiple discharges are supplied in FAST. In the second
stage, when regular sintering takes place, the current is DC
at a level dependent on the powder type. The conductive
powders are heated mainly due to the Joule effect. The
pulsed current may be applied prior to or throughout the
Joule heating cycle. For non-conductive powders heating
occurs through heat transfer from the die and plungers. In
this case, the die and punches are heated through their own
resistance.
Conventional sintering of Ti and Ti alloy powders
requires maintaining a high sintering temperature
(1200–1400 1C) in high vacuum (4104 Pa) for a long time
(24–48 h) [106]. This difficult sintering process limits the use
of sintered Ti and Ti alloys. The aforementioned methods
are useful in that they can easily sinter Ti and Ti alloy
powders, because ionisation in the plasma created by the
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high current discharge can melt the local oxide surface film
on the particles, bringing the particles into contact with
each other, and allowing junctions to be formed. Recently,
EDC has been used to produce commercially pure porous
titanium implants [103,104,107]. Surface analysis performed by Lee et al. indicates that implants produced by
this method allow rapid osseointegration [108]. Lifland
et al. produced porous-surfaced Ti–6Al–4V implants using
the same method. They found the compacts to have yield
strengths ranging from 270 to 530 MPa and ultimate
compressive strengths to range from 390 to 600 MPa
[108]. Using SPS, Kon et al. produced porous Ti–6Al–4V
with a porosity of 32% and compressive strength of
125 MPa [109].
5.3.2. Rapid prototyping
Recently a novel manufacturing process in the form of
three-dimensional printing (3DP) has been used to create
porous implants with controlled size, shape and distribution of the porosity [110,111]. 3DP is a rapid prototyping
(RP) technology, used to create complex three-dimensional
parts directly from a computer model of the part, with no
need for tooling. The sequence of operations is depicted in
Fig. 15.
Parts created using 3DP is a layered printing process
where the information for each layer is obtained by
applying a slicing algorithm to the computer model of
Fig. 15. The use of the 3DP process can be beneficial in the fabrication of prototypes. This method can produce high accuracy filler structures for the
fabrication of porous metals.
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the part. Parts are created inside a cavity that contains a
powder bed supported by the moving piston. Each new
layer is fabricated through lowering of the piston by a layer
thickness and filling the resulting gap with a thin
distribution of powder. Similar to ink-jet printing technology, a binder material selectively joins powder particles at
sites where they have to be welded. The layering process is
repeated until the part is completed. Following a heat
treatment, which consolidates the bonded material, the
unbound powder is removed, leaving the fabricated part
behind.
In a variation of the standard lost-wax casting process,
3DP was applied to produce ceramic shells without the
need of a wax pattern, directly from CAD representation
[111]. Molten metal was infiltrated into sub-millimetre
mould cavities to form the desired parts, and the mould
material was subsequently removed to produce a functional, textured metal part. A broad range of textures was
produced to demonstrate the capabilities of the 3-D
printing process as shown in Fig. 16, where the CAD
renderings of unit cell surface textures are shown. This unit
cell is built from cubic primitives, which are combined in
Fig. 16. Computer-aided design files and scanning electron micrographs of different macro-textured surfaces as built with the help of 3DP. Reprinted with
permission from [110].
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various patterns to produce the desired surface. Each cube
on the CAD model represents the minimum feasible feature
size achieved from the individual 3DP mould. This unit cell
texture can be duplicated over a part surface to create a
uniform texture. As shown on the right-hand side of figure,
such surface textures were successfully cast by a reversegravity process using 3-D printed moulds made of Co–Cr
alloys.
6. Characteristics of porous metals that affect their
performance
The fixation efficacy of coatings is strongly dependent on
the geometric characteristics of the coating layer. The
different fabrication techniques that have been reviewed
will form porous materials with varying pore shapes and
size ranges. In general, no significant difference in biologic
response as a function of pore shape has been reported
[112]. However, it is recognised that one of the critical
factors for bone ingrowth is the size of interconnecting
pores and several investigators have studies bone ingrowth
into porous systems in this respect [59,63,113–121].
Although optimum pore size required for implant fixation
remains undefined, the consensus is that in order to
optimise mineralised bone ingrowth, pore sizes between
100 and 400 mm are necessary [113]. However, Bobyn et al.
[120] showed effective bone ingrowth into porous coatings
with pore sizes down to 50 mm, and Itälä et al. [121]
demonstrated the formation of an osteonal bone structure
in pore sizes as small as 50 mm. Conversely, when the pore
size is increased beyond 1 mm there seems to be an
increasing tendency for the formation of fibrous tissue
[122]. The porosity of most implants is usually determined
to compromise between maintaining the mechanical
strength of the implant while still providing adequate pore
size for tissue ingrowth.
Several investigators have penned terms to describe the
biological events in porous implant stabilisation. The term
osteoconduction refers to the situation where bone can grow
on the surface of an implant [123]. This phenomenon is
frequently seen in the case of bone implants. Materials of
low in vivo acceptance, such as silver or copper show little
or no osteoconduction. The term osseointegration was
originally described by Branemark et al. [124], as the
intimate contact of bone tissue with the surface of a
titanium implant, but it is now generally accepted as the
rigid fixation of an implant by the formation of bony tissue
around the implant without the growth of fibrous tissue at
the bone implant interface [123]. Osseointegration is
dependent on osteoconduction but whilst the osteoconductive response may be short lived, successful osseointegration maintains its bone anchorage over extended periods.
The term bone ingrowth refers specifically to bone formation within a porous surface structure [24]. Bone ingrowth
requires osteoconductive surfaces and will lead to successful osseointegration of the implant.
2665
The physiologic response to an inserted bulk or porouscoated implant is comparable to the healing cascade of
cancellous defects, with newly formed tissues infiltrating
the void spaces of the porous material. Capillaries,
perivascular tissues, and osteoprogenitor cells migrate into
porous spaces and incorporate the porous structure by
forming new bone. Similar to primary fracture healing with
stable osteosynthesis, no intermediate fibro-cartilaginous
stage occurs. With initial sufficient stability, the early tissue
that infiltrates the pores, differentiates to bone by either (1)
direct bone formation within the pores, or, (2) appositional
bone growth from the adjacent bone into the porous region
[125]. Initial implant movement relative to host bone can
result in attachment by a non-mineralised fibrous connective tissue layer, inhibiting bone formation within, or
ingrowth into the pores [126]. A quantitative description of
the maximum limit of movement to allow bone ingrowth
has not been determined although observations reported
by Pilliar et al. suggest that, if shear displacements at the
bone-implant interface are greater than about 30 mm, bone
ingrowth will be inhibited [127]. This report and earlier
studies indicate that greater initial implant movement
nevertheless could result in implant fixation but by a nonmineralised fibrous connective tissue [128]. Consequently,
the initial stability achieved at the implant bone interface,
following implant replacement, is a strong determinant of
the type of biologic fixation, if any that eventually
develops.
The lack of direct contact at the porous surface-bone
interface also has a negative effect on bone ingrowth and
the eventual strength of fixation of the implant. This
apposition is often not achieved because of the design of
the implant or instruments as well as the operative
technique. Several investigators have performed controlled
gap models to demonstrate the inhibiting effect of gap size
[113,129,130]. A study by Dalton et al. suggests that gaps
greater than 1 millimetre significantly affect the attachment
and bone ingrowth [131]. Clinically, these studies imply the
significance of accomplishing direct implant and bone
contact by the surgical technique. Implant related factors,
such as material characteristics, design characteristics,
porosity characteristics, and the use of growth stimulators
also heavily influence the eventual outcome of the implant
fixation.
The selection of materials for medical applications is
usually based on considerations of in vivo performance and
physical functionality. When metals and alloys are
considered, the susceptibility of the material to corrosion
and its local and systemic effects are central aspects for
consideration. Corrosion resistance of the currently used
316L stainless steel, Co–Cr, and titanium-based implant
alloys relies on their passivation by a thin surface layer of
oxide. Of the alloys accepted for conventional (smoothsurfaced) surgical implant fabrication, only porous-coated
316L stainless steel implants may have compromised
corrosive resistance, due to potentially higher corrosion
rates that can occur in vivo with the irregular and porous
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surface geometry. Stainless steel (316L) forms a less stable
oxide layer and is more susceptible to crevice corrosion
than are the other currently used implant alloys [43].
Commercially pure titanium has proven its suitability as an
implant material in bone surgery over many years in the
fields of osteosynthesis, oral implantology, and joint
prosthetics [132,133]. Furthermore, Head et al. describe
titanium as the material of choice for cementless femoral
components in total hip arthroplasties [3]. Sotereanos et al.
suggested the use of Co–Cr alloys as the material of choice
based on a study of porous-coated femoral prostheses
[134]. Finally, a study by Galante et al. that compared Ti
and Co–Cr fibre-made porous coatings indicated no
significant differences for bone ingrowth rates with these
two different materials [65].
Recently, porous Ni–Ti inter-metallics alloys have
received increasing attention as potential materials for
porous orthopaedic implants, due to their good corrosion
resistance and unique mechanical properties [64,82,135,
136]. Ni–Ti is comparable to 316L stainless steel, in that
both share the presence of an adherent TiO2 surface oxide,
which in the case of Ni–Ti prevents nickel dissolution and
release [137]. Ni–Ti alloys of near equiatomic parts of
nickel and titanium, known as Nitinol, exhibit three
properties not commonly observed in metallic materials:
thermal shape memory, superelasticity, and high damping
properties [137,138]. NiTi can be strained several times
more than ordinary metal alloys without being plastically
deformed. It provides a possibility for making selfexpanding and self-locking implants [139].
Although several studies have shown excellent bone
ingrowth into porous metallic systems, there have been a
number of cases in which porous surfaced prostheses have
been retrieved with minimal or no bone ingrowth present
[140,141]. Even though metals such as titanium and Co–Cr
are bio-inert, they do not bond directly to bone. A fibrous
layer intervenes between the implant and bone. Bioactive
materials are designed to induce a specific biological
activity, which can lead to strong bonding to bone [142].
The essential requirement for an artificial material to bond
to living bone is the formation of a biologically active
bone-like apatite layer on its surface in the living body
[143]. Ducheyne et al. reported increased bone ingrowth
rates, due to thin hydroxyapatite coatings applied to fibremade, porous-surfaced 316L implants [144]. Later studies
of implants with thin, plasma-sprayed hydroxyapatite
coatings over powder-made, porous Ti surfaces have
confirmed the effect of increased rates of bone ingrowth
due to the hydroxyapatite coating [145]. Kokubo et al.
discovered that an in vitro chemical-deposited bone-like
apatite on commercially pure Ti could be induced by an
alkali and heat treatment process followed by soaking in
simulated body fluid (SBF) [146]. A number of investigators have used this approach to produce apatite layers on
porous titanium surfaces [147,148]. A recent study by
Fujibayashi et al. demonstrated extensive bone ingrowth
into such a structure [149].
Finally, another aspect that has been considered in one
class of porous materials, namely the porous surfaced
implants that have been fabricated using plasma spraying
techniques, involves the study of the effect that periprosthetic strain conditions have on the local bone
formation [150]. Thereby, using finite element models it
was predicted that, based on the implant geometry,
appositional bone formation would occur when strain
was less than 8%, whereas new bone formation occurred
when distortional tissue strains were less than 3%. This led
the researchers to develop a mechanoregulatory model that
could allow the prediction of bone ingrowth potential of
porous surfaced implants with various surface geometries.
7. Discussion and future work
The interest in using porous materials for orthopaedic
reconstructive surgery as a means of replacing autografts is
of increasing interest and the large number of scientific
reports confirm this trend. For load-bearing orthopaedic
applications, metals have so far shown the greatest
potential as the basis for such scaffolds, owing to their
excellent mechanical strength and resilience when compared to alternative biomaterials, such as polymers and
ceramics. The focus thereby has mainly been on applications that involve bone ingrowth into the porous scaffolds—either as part of a coating or as a complete matrix.
This has led to the majority of research interest to be drawn
to the development of open-cell porous metals, although
arguably, great potential lies within the use of closed-cell
porous metals, too. In such cases, bone ingrowth would not
be the major interest, but rather the reduction of material
stiffness that has been linked to early implant loosening
following processes of bone loss due to stress shielding [17].
Closed-cell porous metals could serve as materials for the
fabrication of implant stems and have either a porouscoated surface to facilitate bone ingrowth onto their
surfaces for stem fixation or have polished or matt solid
surfaces that could be used with regular bone cement for
their fixation in the bone matrix. The successful employment of both open-cell and closed-cell porous metals relies
on the same requirement that is a suitable fabrication
method that can ensure homogeneously distributed pores
of similar size and shape and cell walls of consistent
thickness and levels of purity and absence of cracks or
crevices that can develop into potential material failure
sites.
It is only now that researchers are starting to understand
the combination of parameters that need to be addressed in
the successful implementation of porous metals in vivo.
This is a multi-factorial design process that has to consider
understanding of material properties, such as corrosion
resistance, passivation levels and potential for bone
adherence; mechanical characteristics including stress–
strain behaviour of the porous metal and its match to
those of bone under various loading conditions [151];
and, finally, parameters involving pore size, shape and
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distribution that will optimise fatigue strength and—in case
of open-cell foams—bone ingrowth.
Although great progress has been made with the various
available fabrication processes in manufacturing both
closed-cell and open-cell porous structures, certain limitations continue to exist. Most current techniques that use
foaming agents, either in solid state sintering processes or
in molten metal techniques have limited control over pore
distributions and densities, and are only capable—at best—
to achieve these over large areas rather than at specific
desired locations in the matrix. Probably, the most
promising technique in controlling local densities is the
CVD technique combined with the use of a replica
structure [92], but there may be issues concerning the
fatigue strength of the resulting hollow struts. Nonetheless,
the results achieved by this technique may be indicative of
the potential to obtain even functionally graded porous
metals in the future, if using adequately shaped polyurethane foam substrates.
Engineering pore distributions to match the mechanical
properties of bone is commonly accepted to be the next
major improvement in the design of open-cell porous
materials [33]. This means that a technique needs to be
available that will allow the precise positioning of pores in
the 3-D matrix, their inter-connectivity, shape and size. In
our research group we have been working towards this aim
and believe that the use of adequate techniques of RP can
provide the technological platform for this precise pore
positioning [152,153]. However, the question remains as to
how to identify the optimal position, shape, density and
size of pores. The answer to this might lie in simply
reconstructing the structure of trabecular bone from the
anatomical site into which the porous metal is to be
implanted, based on the use of m-CT images [151].
However, this method still produces a mismatch of stiffness
values, because the matrix materials (metals) differ
substantially in mechanical properties compared to those
of bone. An alternative method of obtaining information
that could be used in engineering the parameters of porous
materials might be the precise mapping of loads that occur
in the inter-vertebral disc during various activities of daily
living and to use that information as a design input
information.
Acknowledgements
We would like to thank the postgraduate fellowship
scheme of the Engineering Faculty at the National
University of Ireland, Galway for funding Mr. Garrett
Ryan.
References
[1] Robertson DM, Pierre L, Chahal R. Preliminary observations of
bone ingrowth into porous materials. J Biomed Mater Res 1976;10:
335–44.
2667
[2] Cameron HU, Macnab I, Pilliar RM. A porous metal system for
joint replacement surgery. Int J Artif Organs 1978;1:104–9.
[3] Head WC, Bauk DJ, Emerson Jr RH. Titanium as the material of
choice for cementless femoral components in total hip arthroplasty.
Clin Orthop 1995:85–90.
[4] Simancik F. Introduction: the strange world of cellular metals. In:
Degischer HP, Kriszt B, editors. Handbook of cellular metals.
Weinheim: Wiley-VCH Verlag; 2002. p. 1–4.
[5] Weber JN, White EW. Carbon-metal graded composites for
permanent osseous attachment of non-porous metals. Mater Res
Bull 1972;7(9):1005–16.
[6] Klawitter JJ, Weinstein AM. The status of porous materials to
obtain direct skeletal attachment by tissue ingrowth. Acta Orthop
Belg 1974;40:755–65.
[7] White EW, Weber JN, Roy DM, Owen EL, Chiroff RT, White RA.
Replamineform porous biomaterials for hard tissue implant
applications. J Biomed Mater Res 1975;9:23–7.
[8] Spector M, Michno MJ, Smarook WH, Kwiatkowski GT. A highmodulus polymer for porous orthopedic implants: biomechanical
compatibility of porous implants. J Biomed Mater Res 1978;12:
665–77.
[9] Klawitter JJ, Bagwell JG, Weinstein AM, Sauer BW. An evaluation
of bone growth into porous high density polyethylene. J Biomed
Mater Res 1976;10:311–23.
[10] Cestero Jr HJ, Salyer KE, Toranto IR. Bone growth into porous
carbon, polyethylene, and polypropylene prostheses. J Biomed
Mater Res 1975;9:1–7.
[11] Homsy CA, Cain TE, Kessler FB, Anderson MS, King JW. Porous
implant systems for prosthesis stabilization. Clin Orthop 1972;89:
220–35.
[12] Sauer BW, Weinstein AM, Klawitter JJ, Hulbert SF, Leonard RB,
Bagwell JG. The role of porous polymeric materials in prosthesis
attachment. J Biomed Mater Res 1974;8:145–53.
[13] Hirschhorn J, McBeath A, Dustoor M. Porous titanium surgical
implant materials. J Biomed Mater Res Symp 1971;2:49–67.
[14] Galante J, Rostoker W, Lueck R, Ray RD. Sintered fiber metal
composites as a basis for attachment of implants to bone. J Bone
Joint Surg Am 1971;53:101–14.
[15] Hahn H, Palich W. Preliminary evaluation of porous metal surfaced
titanium for orthopedic implants. J Biomed Mater Res 1970;4:
571–7.
[16] Karagienes M. Porous metals as a hard tissue substitute. I.
Biomedical aspects. Biomater Med Dev Artif Organs 1973:171–81.
[17] Kroger H, Venesmaa P, Jurvelin J, Miettinen H, Suomalainen O,
Alhava E. Bone density at the proximal femur after total hip
arthroplasty. Clin Orthop Relat Res 1998;352:66–74.
[18] Bobyn JD, Glassman AH, Goto H, Krygier JJ, Miller JE, Brooks
CE. The effect of stem stiffness on femoral bone resorption after
canine porous-coated total hip arthroplasty. Clin Orthop Relat Res
1990:196–213.
[19] Bobyn JD, Mortimer ES, Glassman AH, Engh CA, Miller JE,
Brooks CE. Producing and avoiding stress shielding. Laboratory
and clinical observations of noncemented total hip arthroplasty.
Clin Orthop 1992:79–96.
[20] Pilliar RM, Cameron HU, Binnington AG, Szivek J, Macnab I.
Bone ingrowth and stress shielding with a porous surface coated
fracture fixation plate. J Biomed Mater Res 1979;13:799–810.
[21] Engh CA, Bobyn JD. Principles, techniques, results, and complications with a porous-coated sintered metal system. Instr Course Lect
1986;35:169–83.
[22] Sychterz CJ, Topoleski LD, Sacco M, Engh Sr CA. Effect of
femoral stiffness on bone remodeling after uncemented arthroplasty.
Clin Orthop Relat Res 2001:218–27.
[23] Otani T, Whiteside LA. Failure of cementless fixation of the femoral
component in total hip arthroplasty. Orthop Clin North Am
1992;23:335–46.
[24] Kienapfel H, Sprey C, Wilke A, Griss P. Implant fixation by bone
ingrowth. J Arthroplasty 1999;14:355–68.
ARTICLE IN PRESS
2668
G. Ryan et al. / Biomaterials 27 (2006) 2651–2670
[25] Bauer TW, Schils J. The pathology of total joint arthroplasty. I.
Mechanisms of implant fixation. Skeletal Radiol 1999;28:423–32.
[26] Galante JO, Jacobs J. Clinical performances of ingrowth surfaces.
Clin Orthop 1992:41–9.
[27] Pilliar RM. Porous surfaced endosseous dental implants: fixation by
bone ingrowth. Univ Tor Dent J 1988;1:10–5.
[28] Ungethum M, Blomer W. Technology of cementless hip endoprosthetics. Orthopade 1987;16:170–84.
[29] Spector M. Historical review of porous-coated implants. J
Arthroplasty 1987;2:163–77.
[30] Pilliar RM. Porous-surfaced metallic implants for orthopedic
applications. J Biomed Mater Res 1987;21:1–33.
[31] Welsh RP, Pilliar RM, Macnab I. Surgical implants. The role of
surface porosity in fixation to bone and acrylic. J Bone Joint Surg
Am 1971;53:963–77.
[32] Freeman MA, Railton GT. Cementless fixation in endoprosthetics.
Orthopade 1987;16:206–19.
[33] Karageorgiou V, Kaplan D. Porosity of 3D biomaterial scaffolds
and osteogenesis. Biomaterials 2005;26:5474–91.
[34] Yue S, Pilliar RM, Weatherly GC. The fatigue strength of porouscoated Ti–6% Al–4%V implant alloy. J Biomed Mater Res 1984;18:
1043–58.
[35] Kohn DH, Ducheyne P. A parametric study of the factors affecting
the fatigue strength of porous coated Ti–6A1–4V implant alloy. J
Biomed Mater Res 1990;24:1483–501.
[36] Crowninshield RD. Mechanical properties of porous metal total hip
prostheses. Instr Course Lect 1986;35:144–8.
[37] Manley MT, Kotzar G, Stern LS, Wilde A. Effects of repetitive
loading on the integrity of porous coatings. Clin Orthop Relat Res
1987:293–302.
[38] Wolfarth D, Ducheyne P. Effect of a change in interfacial geometry
on the fatigue strength of porous-coated Ti–6A1–4V. J Biomed
Mater Res 1994;28:417–25.
[39] Pilliar RM. Porous biomaterials. In: Williams D, editor, Concise
encyclopedia of medical and dental materials. New York: Pergamon
Press; Cambridge, MA: The MIT Press; 1990. p. 312–9.
[40] Cook SD, Thongpreda N, Anderson RC, Haddad Jr RJ. The effect
of post-sintering heat treatments on the fatigue properties of porous
coated Ti–6Al–4V alloy. J Biomed Mater Res 1988;22:287–302.
[41] Jacobs JJ, Skipor AK, Patterson LM, Hallab NJ, Paprosky WG,
Black J, et al. Metal release in patients who have had a primary total
hip arthroplasty. A prospective, controlled, longitudinal study. J
Bone Joint Surg Am 1998;80:1447–58.
[42] Jacobs JJ, Gilbert JL, Urban RM. Corrosion of metal orthopaedic
implants. J Bone Joint Surg Am 1998;80:268–82.
[43] Ducheyne P. In vitro corrosion study of porous metal fibre coatings
for bone ingrowth. Biomaterials 1983;4:185–91.
[44] Reclaru L, Eschler PY, Lerf R, Blatter A. Electrochemical corrosion
and metal ion release from Co–Cr–Mo prosthesis with titanium
plasma spray coating. Biomaterials 2005;26:4747–56.
[45] Black J. Systemic effects of biomaterials. Biomaterials 1984;5:
11–8.
[46] Becker BS, Bolton JD. Corrosion behaviour and mechanical
properties of functionally gradient materials developed for possible
hard-tissue applications. J Mater Sci Mater Med 1997;8:793–7.
[47] Seah KHW, Thampuran R, Teoh SH. The influence of pore
morphology on corrosion. Corrosion Sci 1998;40:547–56.
[48] Bundy KJ, Williams CJ, Luedemann RE. Stress-enhanced ion
release—the effect of static loading. Biomaterials 1991;12:627–39.
[49] Körner C, Singer RF. Foaming processes for aluminium. In:
Degischer HP, Kriszt B, editors. Handbook of cellular metals.
Weinheim: Wiley-VCH Verlag; 2002. p. 8–14.
[50] Banhart J. Properties and applications of cast aluminum sponges.
Adv Eng Mater 2000;2(4):188–91.
[51] Banhart J, Baumeister J. Deformation characteristics of metal
foams. J Mater Sci 1998;33:1431–40.
[52] Körner C, Singer R. Processing of metal foams—challenges and
opportunities. Adv Eng Mater 2000;2:159–65.
[53] Banhart J. Manufacture, characterisation and application of cellular
metals and metal foams. Progr Mater Sci 2001;46:559–632.
[54] Salito A., Van Osten K.-U., Breme F. Schonende Beschichtungstechnik. Sulzer Technical Review 1998.
[55] Thieme M, Wieters KP, Bergner F, Scharnweber D, Worch H,
Ndop J, et al. Titanium powder sintering for preparation of a
porous functionally graded material destined for orthopaedic
implants. J Mater Sci Mater Med 2001;12:225–31.
[56] Yang YZ, Tian JM, Tian JT, Chen ZQ, Deng XJ, Zhang DH.
Preparation of graded porous titanium coatings on titanium implant
materials by plasma spraying. J Biomed Mater Res 2000;52:333–7.
[57] Rausch G, Banhart J. Making cellular metals from metals other
than aluminum. In: Degischer HP, Kriszt B, editors. Handbook of
cellular metals. Weinheim: Wiley-VCH Verlag; 2002. p. 21–8.
[58] Taylor N, Dunana DC, Mortensen A. Initial stage hot pressing of
monosized Ti and 90% Ti-10% TiC powders. Acta Metall Mater
1993;41:955–65.
[59] Cook SD, Walsh KA, Haddad Jr RJ. Interface mechanics and bone
growth into porous Co–Cr–Mo alloy implants. Clin Orthop
1985:271–80.
[60] Dabrowski JR. Use of powder metallurgy for development of
implants of Co–Cr–Mo alloy powder. Biomed Tech (Berlin) 2001;
46:106–8.
[61] Asaoka K, Kuwayama N, Okuno O, Miura I. Mechanical properties and biomechanical compatibility of porous titanium for dental
implants. J Biomed Mater Res 1985;19:699–713.
[62] Oh IH, Nomura N, Masahashi N, Hanada S. Mechanical properties
of porous titanium compacts prepared by powder sintering. Scripta
Mater 2003;49:1197–202.
[63] Clemow AJ, Weinstein AM, Klawitter JJ, Koeneman J, Anderson J.
Interface mechanics of porous titanium implants. J Biomed Mater
Res 1981;15:73–82.
[64] Rhalmi S, Odin M, Assad M, Tabrizian M, Rivard CH, Yahia LH.
Hard, soft tissue and in vitro cell response to porous nickel–titanium: a biocompatibility evaluation. Biomed Mater Eng 1999;9:
151–62.
[65] Galante J, Rostoker W. Fiber metal composites in the fixation of
skeletal prosthesis. J Biomed Mater Res 1973;7:43–61.
[66] Ronningen H, Solheim LF, Langeland N. Invasion of bone into
porous fiber metal implants in cats. Acta Orthop Scand
1984;55:352–8.
[67] Weiss MB. Titanium fiber-mesh metal implant. J Oral Implantol
1986;12:498–507.
[68] Hoshijima K, Yamamoto H, Yamashita H. Experimental studies of
titanium fiber metal implant for spine fusion. Nippon Seikeigeka
Gakkai Zasshi 1988;62:399–413.
[69] Martell JM, Pierson III RH, Jacobs JJ, Rosenberg AG, Maley M,
Galante JO. Primary total hip reconstruction with a titanium fibercoated prosthesis inserted without cement. J Bone Joint Surg Am
1993;75:554–71.
[70] Bram M. High-porosity titanium, stainless steel, and superalloy
parts. Adv Eng Mater 2000;2:196–9.
[71] Wen CE, Yamanda Y, Shimojima K, Chino Y, Asahina T, Mabuchi
M. Fabrication and characterization of autogenous titanium foams.
Eur Cells Mater 2001;1:61–2.
[72] Wen CE, Mabuchi M, Yamanda Y, Shimojima K, Chino Y,
Asahina T. Processing of biocompatible porous Ti and Mg. Scripta
Mater 2001;45:1147–53.
[73] Tuchinskiy L, Loutfy R. Titanium foams for medical applications.
ASM Conference on Materials and Processes for Medical Devices,
Anaheim, CA, 2003.
[74] Li JP, Li SH, de Groot K, Layrolle P. Preparation and
characterization of porous titanium. Key Eng Mater 2002;218:51–4.
[75] Li JP, Li SH, de Groot K, Layrolle P. Improvement of porous
titanium with thicker struts. Key Eng Mater 2003;240:547–50.
[76] Zhang X, Ayers RA, Thorne K, Moore JJ, Schowengerdt F.
Combustion synthesis of porous materials for bone replacement.
Biomed Sci Instrum 2001;37:463–8.
ARTICLE IN PRESS
G. Ryan et al. / Biomaterials 27 (2006) 2651–2670
[77] Li BY, Rong LJ, Li YY, Gjunter VE. Synthesis of porous Ni–Ti
shape-memory alloys by self-propagating high-temperature synthesis: reaction mechanism and anisotropy in pore structure. Acta
Mater 2000;48:3895–904.
[78] Yea CL, Sung WY. Synthesis of NiTi intermetallics by selfpropagating combustion. J Alloys Compds 2004;376:79–88.
[79] Chu CL, Chung CY, Lin PH, Wang SD. Fabrication of porous NiTi
shape memory alloy for hard tissue implants by combustion
synthesis. Mater Sci Eng 2004;A366:114–9.
[80] Li YH, Rong LJ, Li YY. Compressive property of porous NiTi alloy
synthesized by combustion synthesis. J Alloys Compds 2002;345:
271–4.
[81] Li BY, Rong LJ, Li YY, Gjunter VE. A recent development in
producing porous NiTi shape memory alloys. Intermetallics 2000;8:
881–4.
[82] Assad M, Jarzem P, Leroux MA, Coillard C, Chernyshov AV,
Charette S, et al. Porous titanium–nickel for intervertebral fusion in
a sheep model: part 1. Histomorphometric and radiological analysis
1. J Biomed Mater Res 2003;64B:107–20.
[83] Ducheyne P, Martens M. Orderly oriented wire meshes as porous
coatings on orthopaedic implants. I: morphology. Clin Mater 1986;
1:59–67.
[84] Markaki AE, Clyne W. Magneto-mechanical stimulation of bone
growth in a bonded array of ferromagnetic fibres. Biomaterials
2004;25:4805–15.
[85] Markaki AE, Clyne TW. Magneto-mechanical bone growth
stimulation by actuation of highly porous ferromagnetic fibre
arrays. SPIE international Symposium of Smart Materials, and
micro-smart systems, Biomedical applications of micro and
nanoengineering II. Australia, Sydney, 2004.
[86] Okuno O, Takada Y, Kinouchi Y, Mizutani H, Yamada H.
Corrosion-resistance of magnetic stainless-steel for magnetic attachment. J Dental Res 1995;74:559.
[87] Paulus JA, Parida GR, Tucker RD, Park JB. Corrosion analysis of
NiCu and PdCo thermal seed alloys used as interstitial hyperthermia
implants. Biomaterials 1997;18:1609–14.
[88] Vrijhoef MMA, Mezger PR, Vanderzel JM, Greener EH. Corrosion
of ferromagnetic-alloys used for magnetic retention of overdentures.
J Dental Res 1987;66:1456–9.
[89] Sivakumar M, Mudali UK, Rajeswari S. Compatibility of ferritic
and duplex stainless-steels as implant materials—in-vitro corrosion
performance. J Mater Sci 1993;28:6081–6.
[90] Adell R, Hansson BO, Branemark PI, Breine U. Intra-osseous
anchorage of dental prostheses. II. Review of clinical approaches.
Scand J Plast Reconstr Surg 1970;4:19–34.
[91] Bobyn JD, Stackpool GJ, Hacking SA, Tanzer M, Krygier JJ.
Characteristics of bone ingrowth and interface mechanics of a new
porous tantalum biomaterial. J Bone Joint Surg Br 1999;81:907–14.
[92] Bobyn J, Hacking S, Chan S, Toh K, Krygier J, Tanzer M.
Characterization of a new porous tantalum biomaterial for
reconstructive orthopaedics. A Scientific Exhibit at the Annual
AAOS, Anahein, CA, 1999.
[93] Gibson L, Ashby M. Cellular solids—structure & properties. New
York: Permagon; 1988.
[94] Heiner A, Poggie R, Brown T. Flexural rigidity of laboratory and
surgical substitutes for human fibular bone grafts. J Musculoskel
Res 1998;2:267–72.
[95] Zardiackas L, McCaskill L, Dillon L, Parsell D, Poggie R.
Structural metallurgical and mechanical evaluation of a porous
tantalum biomaterial. Transactions of the 25th Annual Meeting of
the Society for Biomaterials, Providence, RI, USA, 1999. p. 455.
[96] Mitchell D, Bogan J, Parsell D, Poggie R, Zardiackas L.
Compressive and bending fatigue of a cellular porous tantalum
biomaterial. Annual Society of Biomaterials Meeting, Providence,
RI, 1999. p. 466.
[97] Zardiackas L, Parsell D, Dillon L, Mitchell D, Nunnery L, Poggie
R. Structure, metallurgy and mechanical properties of a porous
tantalum foam. J Biomed Mater Res 2001;58:180–7.
2669
[98] Zou X, Li H, Bunger M, Egund N, Lind M, Bunger C. Bone
ingrowth characteristics of porous tantalum and carbon fiber
interbody devices: an experimental study in pigs. Spine J 2004;
4:99–105.
[99] Groza JR, Zavaliangos A. Sintering activation by external electrical
field. Mater Sci Eng A—Struct Mater Properties Microstructure
Process 2000;287(2):171–7.
[100] Miyao R, et al. Fabrication of functionally graded implants by
spark plasma sintering and their properties. J Japan Soc Powder
Powder Metall (Japan) 2000;47(11):1239–42.
[101] Watari F, et al. Effect of spark plasma sintering pressure on the
properties of functionally graded implant and its biocompatibility.
J Japan Soc Powder Powder Metall (Japan) 2002;49(12):1063–9.
[102] Schneider JA, Mishra RS, Mukherjee AK. Plasma activated
sintering of ceramic materials. Ceramic Trans 1996;79:143–51.
[103] Okazaki K, et al. Physical characteristics of Ti–6A1–4V implants
fabricated by electrodischarge compaction. J Biomed Mater Res
1991;25(12):1417–29.
[104] Qiu J, et al. Composite titanium dental implant fabricated by
electro-discharge compaction. Biomaterials 1997;18(2):153–60.
[105] Lifland MI, Okazaki K. Properties of titanium dental implants
produced by electro-discharge compaction. Clin Mater 1994;17(4):
203–9.
[106] Asaoka K, et al. Mechanical properties and biomechanical
compatibility of porous titanium for dental implants. J Biomed
Mater Res 1985;19(6):699–713.
[107] Lifland MI, Kim DK, Okazaki K. Mechanical properties of a
Ti–6A1–4V dental implant produced by electro-discharge compaction. Clin Mater 1993;14(1):13–9.
[108] Lee W, et al. Surface characteristics of a porous-surfaced Ti–6Al–4V
implant fabricated by electro-discharge-compaction. J Mater Sci
2000;35(3):593–8.
[109] Kon M, Hirakata LM, Asaoka K. Porous Ti–6Al–4V alloy
fabricated by spark plasma sintering for biomimetic surface
modification. J Biomed Mater Res B Appl Biomater 2004;68(1):
88–93.
[110] Melican M, Zimmerman M, Dhillon M, Ponnambalam A,
Curodeau A, Parsons J. Three-dimensional printing and porous
metallic surfaces: a new orthopedic application. J Biomed Mater Res
2001;55:194–202.
[111] Curodeau A, Sachs E, Caldarise S. Design and fabrication of cast
orthopedic implants with freeform surface textures from 3-D printed
ceramic shell. J Biomed Mater Res 2000;53:525–35.
[112] Turner TM, Sumner DR, Urban RM, Rivero DP, Galante JO. A
comparative study of porous coatings in a weight-bearing total hiparthroplasty model. J Bone Joint Surg Am 1986;68:1396–409.
[113] Cameron HU, Pilliar RM, Macnab I. The rate of bone ingrowth
into porous metal. J Biomed Mater Res 1976;10:295–302.
[114] Hofmann AA, Bloebaum RD, Bachus KN. Progression of human
bone ingrowth into porous-coated implants. Rate of bone ingrowth
in humans. Acta Orthop Scand 1997;68:161–6.
[115] Bloebaum RD, Bachus KN, Momberger NG, Hofmann AA.
Mineral apposition rates of human cancellous bone at the interface of porous coated implants. J Biomed Mater Res 1994;28:
537–44.
[116] Schliephake H, Neukam FW, Klosa D. Influence of pore dimensions on bone ingrowth into porous hydroxylapatite blocks used as
bone graft substitutes. A histometric study. Int J Oral Maxillofac
Surg 1991;20:53–8.
[117] Harris WH, Jasty M. Bone ingrowth into porous coated canine
acetabular replacements: the effect of pore size, apposition, and
dislocation. Hip 1985:214–34.
[118] Bobyn JD, Cameron HU, Abdulla D, Pilliar RM, Weatherly GC.
Biologic fixation and bone modeling with an unconstrained canine
total knee prosthesis. Clin Orthop 1982:301–12.
[119] Martens M, Ducheyne P, De Meester P, Mulier JC. Skeletal fixation
of implants by bone ingrowth into surface pores. Arch Orthop
Trauma Surg 1980;97:111–6.
ARTICLE IN PRESS
2670
G. Ryan et al. / Biomaterials 27 (2006) 2651–2670
[120] Bobyn J, Pilliar RM, Cameron H, Weatherly G. The optimum pore
size for the fixation of porous surfaced metal implants by ingrowth
of bone. Clin Orthop 1980;150:263–70.
[121] Itälä AI, Ylanen HO, Ekholm C, Karlsson KH, Aro HT. Pore
diameter of more than 100 micron is not requisite for bone ingrowth
in rabbits. J Biomed Mater Res 2001;58:679–83.
[122] Bobyn J, Miller J. Features of biologically fixed devices. In: Simon
S, editor. Orthopaedic basic science. Chicago: American Academy
of Orthopaedic Surgeons; 1994.
[123] Albrektsson T, Johansson C. Osteoinduction, osteoconduction and
osseointegration. Eur Spine J 2001;10:S96–S101.
[124] Branemark PI. Osseointegration and its experimental background.
J Prosthet Dent 1983;50:399–410.
[125] Spector M. Bone ingrowth into porous polymers. In: Williams D,
editor. Biocompatibility of orthopedic implants. Boca Raton: CRC
Press; 1982. p. 55.
[126] Cameron H, Pilliar RM, Macnab I. The effect of movement on the
bonding of porous metal to bone. J Biomed Mater Res 1973;7:301.
[127] Pilliar R, Lee J, Maniatopoulos C. Observations on the effect of
movement on bone ingrowth into porous-surfaced implants. Clin
Orthop 1986;208:108–13.
[128] Pilliar RM, Cameron HU, Welsh RP, Binnington AG. Radiographic and morphologic studies of load-bearing porous-surfaced
structured implants. Clin Orthop 1981:249–57.
[129] Bobyn JD, Pilliar RM, Cameron HU, Weatherly GC. Osteogenic
phenomena across endosteal bone-implant spaces with porous
surfaced intramedullary implants. Acta Orthop Scand 1981;52:
145–53.
[130] Sandborn PM, Cook SD, Spires WP, Kester MA. Tissue response to
porous-coated implants lacking initial bone apposition. J Arthroplasty 1988;3:337–46.
[131] Dalton JE, Cook SD, Thomas KA, Kay JF. The effect of operative
fit and hydroxyapatite coating on the mechanical and biological
response to porous implants. J Bone Joint Surg Am 1995;77:
97–110.
[132] Pohler OE. Unalloyed titanium for implants in bone surgery. Injury
2000;31(Suppl 4):7–13.
[133] Hirschhorn J, McBeath A, Dustoor M. Porous titanium surgical
implant materials. J Biomed Mater Res Symp 1971;1971:49–67.
[134] Sotereanos NG, Engh CA, Glassman AH, Macalino GE, Engh Jr
CA. Cementless femoral components should be made from cobalt
chrome. Clin Orthop 1995:146–53.
[135] Assad M, Chernyshov A, Leroux MA, Rivard CH. A new porous
titanium–nickel alloy: part 2. Sensitization, irritation and acute
systemic toxicity evaluation. Biomed Mater Eng 2002;12:339–46.
[136] Ryhanen J, Kallioinen M, Tuukkanen J, Junila J, Niemela E,
Sandvik P, et al. In vivo biocompatibility evaluation of nickel–titanium shape memory metal alloy: muscle and perineural tissue
responses and encapsule membrane thickness. J Biomed Mater Res
1998;41:481–8.
[137] Medical Applications for Shape-memory Alloys (SMA). IMechE
Seminar. Professional Engineering Publications; 1999. p. 65.
[138] Castleman LS, Motzkin SM, Alicandri FP, Bonawit VL. Biocompatibility of nitinol alloy as an implant material. J Biomed Mater
Res 1976;10:695–731.
[139] Drugacz J, Lekston Z, Morawiec H, Januszewski K. Use of TiNiCo
shape-memory clamps in the surgical treatment of mandibular
fractures. J Oral Maxillofac Surg 1995;53:665–71 (discussion 672).
[140] Collier JP, Mayor MB, Chae JC, Surprenant VA, Surprenant HP,
Dauphinais LA. Macroscopic and microscopic evidence of prosthetic fixation with porous-coated materials. Clin Orthop 1988:173–80.
[141] Cook SD, Barrack RL, Thomas KA, Haddad Jr RJ. Quantitative
histologic analysis of tissue growth into porous total knee
components. J Arthroplasty 1989;4(Suppl):S33–43.
[142] Rawlings RD. Bioactive glasses and glass-ceramics. Clin Mater
1993;14:155–79.
[143] Kokubo T. Bioactive glass ceramics: properties and applications.
Biomaterials 1991;12:155–63.
[144] Ducheyne P, Hench LL, Kagan II A, Martens M, Bursens A, Mulier
JC. Effect of hydroxyapatite impregnation on skeletal bonding of
porous coated implants. J Biomed Mater Res 1980;14:225–37.
[145] Cook SD, Thomas KA, Dalton JE, Volkman TK, Whitecloud III
TS, Kay JF. Hydroxylapatite coating of porous implants improves
bone ingrowth and interface attachment strength. J Biomed Mater
Res 1992;26:989–1001.
[146] Kim HM, Miyaji F, Kokubo T, Nakamura T. Preparation of
bioactive Ti and its alloys via simple chemical surface treatment.
J Biomed Mater Res 1996;32:409–17.
[147] Nishiguchi S, Kato H, Neo M, Oka M, Kim HM, Kokubo T, et al.
Alkali- and heat-treated porous titanium for orthopedic implants.
J Biomed Mater Res 2001;54:198–208.
[148] Kim HM, Kokubo T, Fujibayashi S, Nishiguchi S, Nakamura T.
Bioactive macroporous titanium surface layer on titanium substrate.
J Biomed Mater Res 2000;52:553–7.
[149] Fujibayashi S, Neo M, Kim HM, Kokubo T, Nakamura T.
Osteoinduction of porous bioactive titanium metal. Biomaterials
2004;25:443–50.
[150] Simmons CA, Meguid SA, Pilliar RM. Mechanical regulation of
localized and appositional bone formation around bone-interfacing
implants. J Biomed Mater Res 2001;55:63–71.
[151] Nazarian A, Stauber M, Müller R. Design and implementation of a
novel mechanical testing system for cellular solids. J Biomed Mater
Res Part B: Appl Biomater 2005;73B:400–11.
[152] Ryan G, Pandit A, Apatsidis DP. Reproducibility of porous
microstructure in homogenous metal foams. Transactions of the
30th Meeting of SFB, Memphis, TN, USA, April 27–30, 2005.
[153] Ryan G, Pandit A, Apatsidis DP. Porous metals for tissue
engineering orthopaedic applications. In: Proceedings of Royal
Society meeting ‘engineered foams and porous materials’, London,
UK, November 1–2, 2004.