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Titanium Implant Osseointegration Problems with Alternate Solutions Using Epoxy/Carbon-Fiber-Reinforced Composite

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The aim of the article is to present recent developments in material research with bisphenyl-polymer/carbon-fiber-reinforced composite that have produced highly influential results toward improving upon current titanium bone implant clinical osseointegration success. Titanium is now the standard intra-oral tooth root/bone implant material with biocompatible interface relationships that confer potential osseointegration. Titanium produces a TiO2 oxide surface layer reactively that can provide chemical bonding through various electron interactions as a possible explanation for biocompatibility. Nevertheless, titanium alloy implants produce corrosion particles and fail by mechanisms generally related to surface interaction on bone to promote an inflammation with fibrous aseptic loosening or infection that can require implant removal. Further, lowered oxygen concentrations from poor vasculature at a foreign metal surface interface promote a build-up of host-cell-related electrons as free radicals and proton acid that can encourage infection and inflammation to greatly influence implant failure. To provide improved osseointegration many different coating processes and alternate polymer matrix composite (PMC) solutions have been considered that supply new designing potential to possibly overcome problems with titanium bone implants. Now for important consideration, PMCs have decisive biofunctional fabrication possibilities while maintaining mechanical properties from addition of high-strengthening varied fiber-reinforcement and complex fillers/additives to include hydroxyapatite or antimicrobial incorporation through thermoset polymers that cure at low temperatures. Topics/issues reviewed in this manuscript include titanium corrosion, implant infection, coatings and the new epoxy/carbon-fiber implant results discussing osseointegration with biocompatibility related to nonpolar molecular attractions with secondary bonding, carbon fiber in vivo properties, electrical semiconductors, stress transfer, additives with low thermal PMC processing and new coating possibilities.
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Titanium Implant Osseointegration Problems with Alternate
Solutions Using Epoxy/Carbon-Fiber-Reinforced Composite
Richard C. Petersen
Restorative Sciences, Biomaterials and Biomedical Engineering, University of Alabama at
Birmingham, SDB 539, 1919 7th Avenue South, Birmingham, AL 35294, USA;
richbme@uab.edu; Tel.: +1-205-934-6898
Abstract
The aim of the article is to present recent developments in material research with bisphenyl-
polymer/carbon-fiber-reinforced composite that have produced highly influential results toward
improving upon current titanium bone implant clinical osseointegration success. Titanium is now
the standard intra-oral tooth root/bone implant material with biocompatible interface relationships
that confer potential osseointegration. Titanium produces a TiO
2
oxide surface layer reactively
that can provide chemical bonding through various electron interactions as a possible explanation
for biocompatibility. Nevertheless, titanium alloy implants produce corrosion particles and fail by
mechanisms generally related to surface interaction on bone to promote an inflammation with
fibrous aseptic loosening or infection that can require implant removal. Further, lowered oxygen
concentrations from poor vasculature at a foreign metal surface interface promote a build-up of
host-cell-related electrons as free radicals and proton acid that can encourage infection and
inflammation to greatly influence implant failure. To provide improved osseointegration many
different coating processes and alternate polymer matrix composite (PMC) solutions have been
considered that supply new designing potential to possibly overcome problems with titanium bone
implants. Now for important consideration, PMCs have decisive biofunctional fabrication
possibilities while maintaining mechanical properties from addition of high-strengthening varied
fiber-reinforcement and complex fillers/additives to include hydroxyapatite or antimicrobial
incorporation through thermoset polymers that cure at low temperatures. Topics/issues reviewed
in this manuscript include titanium corrosion, implant infection, coatings and the new epoxy/
carbon-fiber implant results discussing osseointegration with biocompatibility related to nonpolar
molecular attractions with secondary bonding, carbon fiber in vivo properties, electrical
semiconductors, stress transfer, additives with low thermal PMC processing and new coating
possibilities.
© 2014 by the authors; licensee MDPI, Basel, Switzerland.
This article is an open access article distributed under the terms and conditions of the Creative Commons Attribution license (http://
creativecommons.org/licenses/by/4.0/)
Conflicts of Interest
The author declares no conflict of interest.
NIH Public Access
Author Manuscript
Metals (Basel). Author manuscript; available in PMC 2015 January 27.
Published in final edited form as:
Metals (Basel). 2014 December ; 4(4): 549–569. doi:10.3390/met4040549.
NIH-PA Author Manuscript NIH-PA Author Manuscript NIH-PA Author Manuscript
Keywords
titanium; composite; bisphenol polymer; carbon fiber; osseointegration; corrosion; infection;
estrogen; microbiocircuit; semiconductor
1. Introduction
Titanium alloys developed in the 1940s for aircraft were made available to orthopedic
surgeons as biomaterials for bone implants approximately at the same time [1] and were also
tested earlier with cat femurs during the late 1930s [2]. Since World War II, the two
dominant titanium alloys have been 98%–99.6% commercially pure titanium (CPTi) and
titanium with 6% aluminum and 4% vanadium (Ti-6Al-4V) alloy [1-4]. CPTi has four
grades of oxygen content from 0.18%–0.40% that increase the yield strength with variable
other small amounts of metal impurities [1-3]. CPTi is generally reserved for dental
applications due to an extremely stable oxide TiO
2
thin surface layer that resists corrosion
under physiologic conditions [1-3] and forms a fine interfacial direct metal to bone contact
as osseointegration [1-4]. Titanium metal has a relatively low modulus for metal [1-3,5].
Subsequent low modulus materials close to bone then reduce problems with “stress
shielding” so that more uniform stress transfer occurs between the implant and bone to
prevent bone resorption from periods with lack of pressure [1,5]. Ti-6Al-4V has been used
for dental implants and although stronger than CPTi, biocompatibility is a concern from
aluminum and vanadium ions released [3]. Ti-6Al-4V has also been used for orthopedic hip
implant stems, but the Ti-alloy is particularly prone to geometrical notch sensitivity with
crack propagation and further wears excessively as the chief concern [1]. Titanium alloys
are also used to repair craniofacial defects caused by trauma, surgical removal of cysts and
tumors, infections, fractures that do not join and congenital or developmental conditions [4].
However, titanium failures occur and appear related to factors that discourage stabilized
bone osseointegration such as trauma from overloading, micromotion and surgical burden
[6] to support inflammation without proper healing and in a small percentage infection next
to exposed metal surface as the final destructive mechanisms for implant loosening [4,7].
Also, the healing response involves serum protein adhesion to the implant that can promote
bacterial attachment to a biomaterial surface [7].
Recent technology moreover supported through aerospace/aeronautical development with
epoxy/carbon-fiber-reinforced composites has demonstrated far-reaching osseointegration
increases when compared to Ti-6Al-4V alloy in animal research [5]. The bisphenol epoxy
backbone structure was developed early in 1936 as the first synthetic estrogen [8] where
estrogen influences are known to produce anabolic stimulating bone formation and
osteoblast differentiation [5,9-12]. Further, fiber-reinforced composite can offer superior
mechanical properties than metals on a strength-to-weight basis for both strength and
modulus [5,13,14]. Occlusal forces interact with titanium implants more harshly than natural
tooth structure because of intimate bone osseointegration contact without a damping
protective periodontal ligament [15,16] where titanium metal cannot adsorb damaging
energy similar to a polymer matrix composite (PMC) [17]. In fact, in vivo animal testing
with extreme loads produced defects lateral to osseointegration between bone and metal
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implant [16,18]. Conversely, in relation to encouraging test results [5] PMCs with carbon
fiber reinforcement can supply densities/modulus much closer to bone [1,2,5] than titanium
[5,14] for improved mechanical deformation providing viscoelastic damping energy
adsorption/dissipation [2,5,17] and healthy stress transfer with tissues/cell membranes [5].
Also, carbon-fiber-reinforced PMC has electrical conductivity/resistivity properties
bordering similarly on bone properties with polymer insulated carbon-fiber conductive
biocircuits [5,19] to support biocompatible physiological relationships [5]. In addition,
thermoset polymer matrix and carbon fiber both offer covalent bonding opportunity to give
strong bone structure support with excellent osseointegration [5]. Further, epoxy/carbon-
fiber-reinforced PMC does not corrode to release Lewis acid-stimulating metal particles that
can initiate an inflammatory response with aseptic bone implant loosening [5]. Finally, low-
thermal polymer-based thermoset processing allows incorporation of minerals and even low-
temperature organic additives for major tissue design-engineering [5].
2. Corrosion
Corrosion is a diffusion interfacial electron-transfer process that occurs on the surface of
metals. Titanium reacts with oxygen electrochemically rapidly in the presence of water to
form a fine oxide layer of TiO
2
that prevents further oxidation [3,20], Equation (1). The
TiO
2
surface layer protects titanium under normal biologic conditions to regenerate if
removed by reactive corrosion equilibrium products as passivation barrier formation and
confers high corrosion resistance [2,3,21]. Titanium can form an oxide layer 10 angstroms
thick in a millisecond and 100 angstroms in a minute [3,22]. In the passivated state, TiO
2
biomaterials generally corrode less than 20 μm/year [22]. TiO
2
as Ti
4+
and O
2−
with even
numbers as the most common oxidation states [23] are considered to provide molecular
interaction similarities to bone [21] possibly by coordination as simple ionic bonds with
analogous even oxidation states through calcium phosphate mineral, Ca
3
(PO
4
)
2
, from
divalent Ca
2+
and O
2−
[23].
(1)
Still, all metal implants are not perfectly passive in a hostile corrosive biological
environment to have some solubility and are subject to metal dissolution with the formation
of metal cations (M
+
) and electrons (e
), Equation (2) [1,3,21]. Aqueous concentrations of
dissolved molecular oxygen in the tissue react and remove electrons to form hydroxyl anion
[1,3,21], Equation (3), that helps drive corrosion through Equation (2) [3]. Further, metal
cations are removed to polarize water forming a Lewis acid, Equation (4) [21,23,24] that can
then accelerate corrosion through Equation (2). Also, with low pH, normal biologic
extracellular chlorine can form hydrochloric acid [21] that may attack titanium [20,22,25]
with undesirable bone responses [22]
(2)
(3)
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(4)
Capillary distance is a measure of lower oxygen concentration or increased acid and lower
pH where zero O
2
concentrations develop at about a 0.2 mm tissue space [26-28]. Resulting
lower oxygen concentrations near the implant surface without an oxygenated blood supply
are unable to satisfy intracellular mitochondrial requirements during energy synthesis to
form water [29,30], Equations (5) and (6).
(5)
(6)
Organelle mitochondria of the cell produce more electrons and also acid during periods of
lower oxygen concentrations [29,30], Figure 1. Subsequent increasing acid that provides
growing hostile conditions with low pH in the biologic chlorine microenvironment adjacent
to the metal implant can then create breakdown conditions of the generally corrosion-
resistant passive TiO
2
oxide layer to reinitiate more corrosion [22]. In addition to metabolic
mitochondrial acid, the pH might become lower from inflammation and infection
particularly if oxygen is blocked.
Different types of common corrosion have been classified for titanium implants. When acid
breaks down the passive TiO
2
oxide layer on a flat surface pitting corrosion occurs
[1,2,21,22]. On the other hand, geometric implant material confinement of acid produces
increased metal dissolution known at crevice corrosion [1,2,21,22]. Friction between the
TiO
2
oxide layer against another surface causes fretting corrosion [1,2,21,22]. When
titanium is in direct contact with a dissimilar metal that is common to both oral and
orthopedic implants galvanic corrosion occurs [1-3,21,22].
Subsequent electrochemical corrosion products from metal implants are thought to be
damaging on local tissues particularly with respect to low intensity electromagnetic fields
that are known to develop by corrosion and can then inhibit osteoblast growth [31]. Aseptic
loosening of implants is thought to occur as a reaction to metal particles from corrosion that
can produce an electric occurrence with electromagnetic field [31] where lower pH next to a
titanium implant needs overall general consideration [20-28]. Titanium particles from
implants are reduced in size by corrosion over time to commonly produce a dark blackened
tissue stain [32]. Titanium particles found in adjacent soft tissue have been known to
produce inflammation, fibrosis and necrotic tissue while infection was found to be a key
reason for implant failure where pain was further noted as a clinical concern [33]. Microbial
influences can also increase corrosion [1]. In terms of inflammation, titanium metal alloy
particle release from implants can result in osteolysis or bone destruction [34]. Alternatively,
after surgical implant placement chronic inflammation that continually heals can eventually
form a fibrous capsule union between the implant and bone that leads to failure [35]. Also,
inflammation appears to be increased at a disproportionate level to mechanical stress by a
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mineralized-type tooth/metal-implant interface solely as hard tissue connection without a
normal fibrous tooth periodontal ligament [36] that forms a damping protective pad
mechanism [15,16].
3. Infection
Implant failure on occasion is related to infection either directly even requiring antibiotic
treatment or implant loosening from bone destruction with potential bacterial colonization
[1,7]. Implant failure from infection is more injurious with greater complications and risks
than aseptic failure [7,37]. The rate for infection with prosthetic hips ranges from
approximately 0.2%–4% depending on the advanced level of surgery and hospital care [1,7]
while infection with fewer less severe complications becomes more of a factor for dental
implants that extend from bone into the oral cavity and other transcutaneous implants [7].
Infection can occur immediately related to implant surgical placement or years later by
hematological transmission from a distant site through the blood from another location
[1,7,37] or a break in the oral mucosa or skin. Further, implants increase the chance for
bacterial infection by presenting a surface without a vascular blood supply and proper
immune response [1]. Many bacteria are acidogenic/acidophillic to produce acid and also
favor acidic growing conditions to metabolize complex organic compounds for a low pH
capable of dissolving hydroxyapatite as enamel and dentin [38-40]. When acids lower the
pH accelerated chemical degradation of polymer hydrocarbons and amines by hydrolysis
occurs at increased rates [41-46] that could increase bacterial survival from organic nutritive
breakdown products acquired through nearby tissue and cells. Also, cured epoxy polymer
that contains different oxygen bonds [13] can be degraded in situ and in vivo [5].
Destructive low pH tissue environments next to metal implants build from metal Lewis acid
corrosion products [21,23,24] while the implant surface prevents proper oxygen supply to
cells for mitochondrial energy synthesis that produces both free radicals from the electron
transport chain and acid from the proton gradient [26-30]. Subsequent rising acidic
environments next to the implant add to chlorine surface interactions with titanium for
increased corrosion [20-22,25]. Further, titanium metal is not known to integrate with soft
tissue and form a seal that occurs with natural teeth to prevent oral bacterial contaminate
leakage into bone. As a result, any disruptions in the implant/bone osseointegrated interface
provide metal surface areas capable of allowing bacterial adhesion with mucopolysaccharide
formation [1,47] and colonization for biofilm formation [1,7,37,47,48]. Also, implant
surface roughness is a factor that improves bacterial adhesion [47]. In fact, implant
properties that enhance osseointegration by protein adsorption also promote bacterial
colonization [7]. Resulting implant biofilms protect bacteria colonies from host immune
responses, antibiotics and allow bacteria to concentrate nutrients [1,7,37,47,48]. Implant
biofilms even transmit along adjacent tissues to promote long-term infection [7]. In addition,
bacterial colonization produces inflammatory responses that interfere with the bone/implant
osseointegration [4]. Most implant infections do not show up in routine cultures because the
biofilm protects bacterial colonies from releasing microbes [1]. However, because even
small amounts of bacteria colonized can disrupt implant osseointegration, cases of aseptic
loosening are being considered as subclinical bacterial contamination [7].
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Loss of osseointegration through peri-implantitis as a destructive inflammation of bone
supported by infection loosens tooth implants [6,36] with similar influences that are
common to chronic adult periodontitis [36]. Numerous bacterial species identified from
failed dental bone implants are analogous to those found with teeth in corresponding clinical
conditions [36]. Frequently threaded implants for tooth/bone implants [36] might impose
extra risk during progressive chronic implant bone loss by interfering with oral hygiene from
difficult to clean inverted surfaces. Deeper titanium/bone implant infections do not have
comparable conditions to clinical periodontitis where bone is resorbed distant from the
periodontal pocket [36]. Because natural teeth have connections through perpendicular
fibers of the periodontal ligament with bone while titanium implants produce parallel fibers
that may not block bacterial penetration as well as teeth, remote bone loss may be a result in
metal implants [36]. Staphylococci are the chief bacteria involved in orthopedic implant
infections and can produce a biofilm after bacterial adhesion [1,37,47,49]. For later stage
extraoral craniofacial implants, infections have most commonly been identified from skin
bacterial species Staphylococcus aureus [4] that are becoming increasingly resistant to
antibiotic treatment [4,37]. Other bacteria generally assist in chronic craniofacial implant
infection with many different bacteria species identified [4].
4. Coatings
Titanium oxide surface layer forms instantly to a depth of 5–10 nm [3,22] in about a minute
and continues to grow up to 200 nm as the reason for implant osseointegration with bone
[49]. The most popular coating process using a plasma sprayed hydroxyapatite (HA) or
Ca
10
(PO
4
)
6
(OH)
2
produces a roughened surface texture that increases surface area to
improve osseointegration bone attachment [3,49]. The mineral phase for bone is
approximately 60% chiefly as HA with traces of other minerals and the remaining being
25% water and 15% organic compounds [1]. Increasing crystalline HA deposition slows
coating release compared to lower HA crystalline deposition [3]. Commercial HA
deposition ranges from 85% crystalline with 15% tricalcium phosphate or Ca
3
(PO
4
)
2
to 97%
crystalline [3]. However, controversy surrounds deposition of HA that shows improved bone
growth next to the implant compared to the titanium metal surface but some studies suggest
HA is detrimental over longer term use [3,49]. The bond for HA with metal is thought to be
unstable and reduced following ion exchange over time with coating dissolution and even
more dissolution of the tricalcium phosphate [3]. Increased failure of HA coatings over
titanium metal is due to inflammation after coating dissolution and delamination [3] that
would show as small defects to possibly protect bacteria hidden in safety for colonization.
Loss of HA appears to reduce physiologic-type acid buffering by phosphate anion that is
helpful under potential harsh lower pH conditions. Further, HA increases bacterial adhesion
[3] while the HA roughened surface promotes bacterial adhesion growth [3,49] all of which
contributes to peri-implantitis [3,49]. Also, modulus for HA osseointegration with adjacent
bone is sufficiently rigid through less favorable energy dissipation to cause tissue reaction
during applied stress at levels where pressure can also interfere with the HA coating
durability [49]. Nitric acid passivates titanium [3] while electrochemical anodization is a
relatively easy, inexpensive surface treatment used to increase surface texture and also
improve the TiO
2
surface with a thicker layer [3,35]. Defects in a metal crystal lattice scatter
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conduction electrons to increase resistivity [14]. Accordingly, the titanium oxide surface
film produced at the anode has been shown to be less conductive with higher resistivity than
the metal titanium [35,50] that may provide new biocompatibility properties for implant
osseointegration [5,50]. In addition, the TiO
2
surface thickness increases with increasing
process temperatures that increases surface roughness, surface energy [50] and hardness [51]
while reducing the contact angle [50] as a measure for increasing surface wetting [1]. No
surface modifications have been found to counteract problems of infection other than
uniform bone/implant osseointegration coverage contacts. Another area of interest that has
shown possibilities for success include studies with bioactive bone morphogenic protein
(BMP) in repairing bone defects to enhance bone growth next to the implant especially since
proteins adsorb onto the implant surface before cell contact [49].
5. Polymer Matrix Composites (PMCs)
5.1. Results for PMC Biocompatibility
Osseointegration and antimicrobial properties are repeatedly hard to realize with titanium/
titanium alloy implants [4], probably because biocompatibility with function is difficult
using metal [52]. Although polymers have been identified for biomaterial use because of
high biologic functionality, polymers lack mechanical strength needed with hard tissue
implants [52]. In terms of polymer biocompatibility with sufficient strength, PMCs using
high-strength fibers provide answers [5]. Fibers are the strongest and possibly the stiffest
forms of a substance matter [53]. When combined into a thermoset cure crosslinking
polymer matrix, fiber-reinforced composite materials provide design possibilities for
ultimate potential in bone implant osseointegration toward biocompatibility with biofunction
[5], Table 1. Most importantly, fiber-reinforced PMCs compete with metals especially on a
strength-to-weight basis in required mechanical properties.
In comparison to a new bisphenol-epoxy/carbon fiber-reinforced composite implant
material, titanium alloy Ti-6Al-4V produces significantly less bone forming near the implant
with much lower levels of osseointegration contact in a bone-marrow animal implant model
[5]. After two weeks, major breakthrough differences were apparent when comparing lateral
cross-sectional percent bone area (PBA) for epoxy/carbon fiber PMC to Ti-6Al-4V alloy
implanted midtibial in vivo using an animal model, Figure 2a,b [5]. At 0.1 mm distance from
the implant PBA increased from 19.3 ± 12.3 with the titanium alloy implant to 77.7 ± 7.0
with the PMC, p < 10
−8
. At 0.8 mm distance PBA increased from 10.5 ± 5.3 with the metal
alloy to 41.6 ± 13.9 with the PMC, p < 10
−4
[5].
Typical histology ground sections for the epoxy carbon fiber PMC and titanium-6Al-4V
alloy as average Histomorphometry PBA measurements are presented in Figure 3a,b.
Osseointegration for the experimental epoxy carbon fiber PMC was broad along the length
of the implant with structural pore-bearing organization for oxygen and nutrient
accessibility. Conversely, titanium-6Al-4V alloy osseointegration was rare and
nonstructured, Figure 4a,b.
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The extent of bone formation for the epoxy/carbon fiber PMC is presented with a horizontal
section to better appreciate the exuberant extent of bone formation inside the bone marrow
that is normally not seen physiologically, Figure 5.
Normal difficult-to-see X-rays show how bone grows through the bone marrow space
alongside the epoxy/carbon fiber PMC implant where bone does not usually grow, Figure
6a,b.
A photograph provides evidence of the strong osteogenic response for the epoxy/carbon
fiber implant with bone growing above the outer cortical bone onto the PMC surfaces,
Figure 7.
5.2. Nonpolar Molecular Attractions with Secondary Bonding
Bisphenol-epoxy/carbon-fiber PMC provides biocompatibility with biologic function
through both the polymer matrix and fiber reinforcement [5]. Epoxy is a thermoset
crosslinking cured polymer and considered polar or more accurately covalently polar in
comparison particularly to nonpolar thermoplastic hydrocarbon-type polymers [13,24].
Because of the presence for possible retained amine, ether or epoxide groups with oxygen
and nitrogen atoms in an epoxy polymer [13] increased polarity is expected for a nonpolar
hydrocarbon [24]. A covalent bond is considered nonpolar when electrons are shared equally
with electrons paired in overlapping orbitals [24]. However, when an electron pair is not
shared equally when linking two atoms, the bond is considered covalent-polar at varying
degrees depending on the nature of electron sharing connecting the two atoms involved [24].
The bond polarity is due to the electronegativity differences involving two separate atoms
[24]. For example, bonds between two carbon atoms are identical and nonpolar while bonds
with a carbon atom and hydrogen atom are basically nonpolar containing similar
electronegativities for both the carbon and hydrogen atoms [24]. On the other hand, bonds
linking carbon with oxygen or nitrogen are polar covalent with larger electronegativities for
oxygen and nitrogen that more strongly attract the bonding pair of electrons [24]. Further,
estrogen factors are present from bisphenol polymers [5,8,9-11,54] with a backbone derived
from one of the first synthetic estrogens [5,8]. Subsequent physiologic actions of estrogen on
bone include skeletal growth, increased osteoblast activity and retained Ca
2+
and HPO
4
2−
mineralization due to organic bone matrix formation [30]. Also, estrogen and a precursor for
resin, bisphenol A, protect the ovary from degeneration, uterine shrinking and bone loss in a
concentration dependent manner [30,54]. Bisphenol A has also shown increased adult rat
femur length without loss of strength [55] and decreased levels of micronuclei in bone
marrow reticuloctyes [56]. In terms of biologic compatible uses, bisphenol A epoxy has
approval level for food contact with coating the inside of food cans to resist corrosion [56]
and in dental composite fillings [3,56].
For a biologic comparison, the cell membrane that comes in contact with a foreign implant
material is composed of lipids, proteins and carbohydrates [30] all of which are similar in
nature to polarity closer to the bisphenol epoxy than a metal. For instance, a cell membrane
is approximately 50:50 lipid:protein by mass weight [30]. The membrane lipids are
amphipathic with a hydrophilic (polar) globular head and hydrophobic (nonpolar) fatty acid
tail [30]. Proteins as hydrocarbons with nitrogen and oxygen amide bonds are found inside
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the membrane and peripherally [30]. Cholesterol is a precursor to estrogen and found in the
membrane to help maintain membrane fluidity [30]. Closed shell molecules attract one
another through van der Waals forces because of the partial charges in polar covalent
chemistry that further includes the small nonpolarity electronegative differences in
hydrocarbons through multipolar effects [57] resulting in an intermesh of related molecular
chains attracting one another. Subsequent similarities in molecular forces of attraction then
exists in variation between the thermoset cure bisphenol polymers with the plasma cell
membrane [5,30] and organic portions of the bone matrix [1,2] as forms of material
biological function [5]. Consequently, bone-marrow precursor cells for the bone-forming
osteoblasts apparently are recruited toward the bisphenol epoxy implant composite by
similar chemical molecular structures to then form mature bone [5]. Regarding stress
transfer, epoxy/carbon-fiber PMC bone plates have been compared with stainless steel and
titanium in human forearm fractures to take advantage of lower modulus material with less
stiffness and better bone response while most of the PMCs produced thin fibrous capsules
grown next to the plates [58].
5.3. Carbon Fiber Biocompatibility
Carbon fibers also appear to stimulate strong cell recruitment during the extensive bone
formation with the bisphenol epoxy implant PMC. Carbon fibers demonstrated extensive
biocompatibility with bone as evidenced from the in vivo bone marrow implant testing
through separate different mechanisms [5]. Carbon fibers are oxidized approximately 20%
as received with R–COOH and R–COH surface groups [59] that should attract many
biologic molecules similarly as hydrocarbons with oxygen through van der Waals forces
[57]. Carbon fiber condensation reactions would provide strong covalent bonds through cell-
membrane lipid fatty acids/phosphate/amino-acid end groups, bone phosphate and some
organic portions of the bone matrix. Although fibers have strongest strengths in tension,
fibers are weak in the transverse direction [14,53]. As a result, carbon fibers were found
broken and in pieces alongside the implant with strong osseointegration bone association
that could have pulled carbon fiber reinforcement sideways in the weak transverse direction,
Figure 8a,b.
Carbon fibers not only stimulate osteoid bone matrix formation, Figure 9a, but further
encourage soft tissue attachments, Figure 9b. In fact, carbon fibers have been tested with
apparent biocompatible success for ligament replacements in human knee reconstruction
demonstrating concentric fibrous layers surrounding a carbon fiber core of mechanically
sound intact fibers [60].
Because normal low oxygen concentrations in bone marrow further produce acids during
mitochondrial energy synthesis, epoxy polymer is softened and pulled away from the
implant by bone attached with carbon fibers, Figure 10a. Small portions of carbon fiber are
eventually degraded into a fine particulate smear layer on the very outer surface immediately
next to the bone. Epoxy polymer is even broken down within the implant itself so that
noncalcified osteoid is evident well into the implant and surrounding individual carbon
fibers for heightened levels of osseointegration, Figure 10b.
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By measure of bisphenol epoxy polymer degradation with depth of bone osseointegration
into the carbon-fiber PMC, a defect in the implant surface can apparently reduce oxygen
concentrations more than elsewhere to lower the pH. The osteocyte bone-forming cell
involved tunnels into small spaces to extend cytoplasmic processes that secrete degrading
enzymes and bone matrix proteins as osteoid [61]. Potential biologic relevant nitric acid
chemistry has previously been considered in prior publications that it attacks bisphenol
aromatic rings supported by a protein enzyme [5]. Figures showing bone to implant
attachments indicate that covalent bonding with the carbon fibers by electron pair sharing is
a chief bond mechanism for osseointegration while polymer covalent bonding appears
possible. Also, mechanical retention develops as polymer degrades for strong bone
ingrowth. On the other hand, titanium electron bonding is ionic with mineralization between
bone and the TiO
2
surface oxide layer.
Carbon fibers are electrically conductive [5,14] and with an insulating polymer coating
become micro-biocircuits in a PMC [5]. Previous description of the implant
microenvironment is found earlier with corrosion that describes the lower oxygen
concentrations. As the distance increases from the blood supply oxygen concentrations
become lower resulting in mitochondrial metabolic production of electrons and acid
[5,26-30]. Subsequent mitochondrial electrons during hypoxia are then able to channel fast
through carbon fibers electrochemically to areas of lower negative charge and lower electron
concentrations [5]. Bone formed cells then preferentially seek carbon fibers to discharge
excess electrons produced from the electron transport chain during mitochondrial energy
synthesis concurrent with hypoxia, otherwise damaging free radicals could be produced [5].
Conductivity confers potential to remove inflammatory surgical free radicals [5] to form
possible covalent bonds with exposed unpaired electrons [62] from the polymer by pH
degradation. Overall, carbon fibers act as a permanent antioxidant to distribute free radicals
that could prevent bone growth [5].
5.4. Electrical Biocompatibility and Semiconducting Properties
Electrical properties of cells have been studied most extensively at the plasma cell
membrane level with a voltage potential of approximate −80 mV but can range from about
−50 mV to −90 mV where the intracellular fluid is more negative with respect to the more
positive extracellular biologic fluid [30]. The plasma cell membrane is composed of fluid
lipid oils also structured intracellularly with protein fibers and extracellularly with divalent
calcium that can form cements as calcium hydroxide and calcium oxide, form secondary
bonds as calcium bicarbonate, produce inorganic mineral apatite as calcium phosphate, and
thin elemental calcium channels [29]. Both protein fibers and fibrilar nanocalcium metals act
as conducting biocircuits with small nanometer diameters to provide efficient electron flow
[29]. Cell nanocircuits are important due to possible excessive electrons that need to be
distributed through electrochemical gradients for uniformity to prevent high concentration
build-ups that follow exponential rates for electron transfer [5,63]. However, unfortunate
high electron current might be excessive and disintegrate small calcium or protein-type
nanocircuits along the outside of the plasma cell membrane. However, semiconducting
cellular materials that appear to exist at the plasma membrane phosphate-head-group/water
interface next to susceptible extracellular nanocircuits [5,64] could safely adsorb and
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conduct excessive electrons until normal undamaging flow is reestablished. Similar use for
semiconductors is well-known for microelectronic circuits that are stacked on top or lie
within a silicon semiconductor wafer with a resistivity of approximately 3000 Ωm [14]. To
better appreciate differences in electric currents that occur between metals that are
conductors, polymer insulators and various semiconductors, resistivities are presented in
Table 2.
In terms of potential problems arising without proper electron distribution, higher-than-
normal electron concentrations can enter into free-radical crosslinking reactions to produce
structured molecules [62]. The molecular structure then has the ability to interfere with
normal biological diffusion or flow to prevent nutritive delivery to cells and even oxygen
can be blocked that complicates physiology into pathological states [62,69]. Electron
transfer reactions are extremely fast [63] and become particularly prevalent when free
radical concentrations build which is the condition during disease with pathology [62,69]
that should require fast conduction unloading of excess cell electrons. Also, high free-radical
concentrations might encompass problems related to surgical inflammation as tissue heals.
By similar free-radical electron transfer chemistry, biologic crosslinking could explain the
coarse or clumping chromatin of DNA to DNA or DNA to protein [69] and protein
agglomeration with insoluble accumulation [70,71] that overall could interfere with implant
healing. Subsequent carbon-fiber-reinforced PMC has electrical conductivity/resistivity
properties bordering on semiconducting bone properties also with polymer insulated carbon-
fiber conductive biocircuits to support vital biocompatible physiological relationships
[2,5,14,19] in preventing electron free-radical build up related to damaging increased
molecular structure [62].
A safer semiconducting biomaterial surface will provide a more physiologic interface for
better biocompatible faster electron transfer interaction with vulnerable nanocircuits of
susceptible cell membranes [5]. As a well-studied relationship, titanium implant
biocompatibility has been emphasized particularly with respect to the corrosion resistant
surface titanium dioxide film. More specifically, TiO
2
surface provides the special property
for implant osseointegration with bone even at an extremely small thickness down in a range
from 5–10 nm to 100–200 nm [49]. Resistivity values for titanium dioxide as a
semiconductor are shown for mineralized rutile in a range of approximately 29–910 Ωm
[65]. Corresponding similar relative semiconducting resistivity magnitudes are found with
bone, a plasma cell membrane phospholipid/water interface model, physiologic saline and
the new highly successful bisphenol-epoxy-polymer/carbon-fiber composite implant
material, Table 2. Therefore, semiconduction apparently plays a role at some level in
biocompatibility for implant osseointegration.
5.5. Stress Transfer
PMCs with carbon fiber reinforcement can supply densities/modulus much closer than
titanium [5,14] to bone [1,2,5] for improved mechanical deformation by viscoelastic
damping energy adsorption/dissipation [2,5,17] and healthy stress transfer with tissues/cell
membranes [5]. Although carbon fibers appeared chemically inert, polymer softening by
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lowered pH created conditions that degraded the polymer with an expected much lower
modulus for far easier deflection when mechanical stress was applied by bone.
5.6. Additives for Low Thermal PMC Processing
Thermal processing of epoxy PMC thermosets can range from room temperature cure up to
less than 200 °C [13,53] compared to far higher temperatures for ceramics or metals [14].
Consequently, additives for epoxy PMC can include inorganic filler or organic compounds
carefully selected for specific implant biocompatible design purposes. In addition to
covalent bonding and polymer softening with bone ingrowth for osseointegration, ionic
bonding mineralization by inorganic fillers as highly stable low-soluble crystalline HA can
be provided. For cell recruiting, phosphate from HA could attract phosphate headgroups
from phospholipid fatty acid plasma cell membranes and further calcium from HA could
attract extracellular plasma cell membrane calcium that cements and mineralizes on the
nanoscale for osseointegration potentials between forming bone and the implant. Also,
phosphate anion acts as a physiologic-like buffer to counteract possible acids produced by
hypoxia particularly next to an implant with inflammatory reactions from surgery. Particle
HA can be infused as normal filler during the PMC resin infusion process with the carbon
fibers. In fact, HA particle filler will be surrounded and retained by polymer completely so
that the implant surface can be polished to a perfect smooth finish to reduce bacterial
adhesion by most of the surface roughness mechanisms. In terms of preventing bacterial
colonization, Triclosan is a highly stable hydrophobic and nonpolar crystalline powder
antimicrobial that will incorporate into resin for PMC infusion [72].
5.7. Biocompatibility Coatings
Resorbable coatings are another feature to consider after the bulk material implant shape is
set where thermal process can be controlled carefully for temperature sensitive proteins.
Highly soluble calcium phosphate is an alternative to HA for rapid release during the
stabilization phase with bone-to-implant osseointegration during healing. Tissue engineering
design principles for bone implant osseointegration developed for long-term bulk-material
application can then be applied for quick release in an outer resorbable coating to enhance
quick implant stabilization with surrounding bone. As examples, low crystalline HA
dissolves faster than highly stable crystalline HA to help speed initial bone growth, estrogen
can enhance nonpolar lipid membrane and other organic attraction forces for improved cell
recruitment, conductive particles or carbon nanotubes can draw in inflammatory free-
radicals with other excess electrons and antimicrobial/antibiotics can be added to control
bacteria introduced during surgical implant insertion.
6. Conclusions
Osseointegration bonding occurs by different covalent electron sharing and ionic
mineralization mechanisms. TiO
2
osseointegration produces ionic bonds by even oxidation
states that act in coordination with the mineralization phase of bone. PMC osseointegration
appears to produce covalent bonds by free-radical crosslinking with exposed unpaired
electrons of the polymer following acid degradation while organic portions of the bone
matrix or bone-cell plasma membrane condense by covalent bonding onto acid or hydroxyl
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groups of the oxidized carbon fibers. Further mechanical interlocking is achieved with
rougher surfaces and with the PMC by acid degradation polymer removal can occur even
with possible bone growth surrounding individual 7 μm diameter carbon fibers. Low pH
polymer softening by acid is considered now to aid in adsorbing excessive stresses by a
protective damping mechanism. Low temperature thermoset polymer cure allows fillers and
organic additives to be incorporated by planned design with new tissue engineering for bone
implants toward biosuccess. Fillers and additives can be included either in the bulk implant
material that is polished to reduce microbial attachment colonization or in extremely mild
resorbable coatings for rapid release to stabilize the initial implant surgical placement.
Future research directions should examine implications clinically for the robust benefits and
also surgical problems particularly during possible revision taking into account such strong
osseointegration for the bisphenol-epoxy/carbon-fiber implant.
Acknowledgments
National Institutes of Health funding Grant No. T32DE14300; Jack E. Lemons, Technical Experimental Protocols,
Director Laboratory Surgical Implant Research and Implant Retrieval Analysis Center Department of Surgery,
School of Medicine, Department of Biomedical Engineering, School of Engineering, Department of
Prosthodontics-Biomaterials, School of Dentistry, University of Alabama at Birmingham; Michael S. Reddy,
Experimental Testing with Biological Technical Advice, Dean, School of Dentistry, University of Alabama at
Birmingham; Michael S. McCracken, Animal Surgery, Associate Dean, School of Dentistry, University of Alabama
at Birmingham, Patricia F. Lott, Ground Section Preparation, Director Center for Metabolic Bone Disease-
Histomorphometry and Molecular Analysis Core Laboratory, University of Alabama at Birmingham, National
Institutes of Health Grant No. P30-AR46031; Preston R. Beck, Technical Advice, PMC Acid Testing and some
Photography, Department of Prosthodontics-Biomaterials, School of Dentistry, University of Alabama at
Birmingham.
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Figure 1.
(a) Mitochondrial electrons combine with protons and molecular oxygen to produce water;
(b) Mitochondria with enzymes involved in ATP energy synthesis depict relationship of
outer membrane to the intermembrane space and inner membrane.
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Figure 2.
Implant percent bone areas comparing epoxy/carbon fiber PMC to Ti-6Al-4V alloy (a)
Distance 0.1 mm from implant (b) Distance 0.8 mm from implant. (error bars ±1 standard
deviation).
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Figure 3.
Lateral cross-sectional toluidine blue stain section at 2× magnification in a rat tibia bone
marrow implant model (a) Epoxy carbon fiber PMC; (b) Titanium-6Al-4V alloy.
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Figure 4.
Toluidine blue stain osseointegration for lateral implant sections at 20× magnifications (a)
Coordinated epoxy/carbon fiber osseointegration (b) Isolated titanium-6Al-4V
osseointegration.
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Figure 5.
Sanderson’s stain epoxy/carbon fiber PMC horizontal section at 2× magnification in the
marrow space shows mature organized pores in osseointegrating bone with the implant.
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Figure 6.
X-rays epoxy/carbon fiber PMC (a) Lateral view (b) Frontal view.
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Figure 7.
Photograph of epoxy/carbon fiber composite extending above tibial cortical bone with bone
stimulated sufficiently to further grow upward along the side of the PMC carbon-fiber
implant.
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Figure 8.
Lateral cross-sectional histology section at 40× magnification by toluidine blue of epoxy/
carbon fiber PMC implant with broken carbon fibers pulled transversely away from the
implant. (a) Carbon fibers are broken and pulled away from implant by bone in the
transverse direction to open up small pore space at the PMC implant surface allowing
minimal oxygen access; (b) After carbon fibers are split and pulled away from the PMC
implant, bone osseointegrates entirely around small carbon fiber segments with a large pore
remaining at the implant surface.
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Figure 9.
Photographs (a) implant extends above cortical bone with exuberant osteoid production
stimulated from small carbon fiber fragments extruded out of the marrow space; (b)
dissected soft tissue overlying the cortical bone integrated with carbon fiber fragments from
the end of the implant.
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Figure 10.
Horizontal cross-sectional histology with Sanderson’s stain at 40× magnification. (a) Bone
osseointegration next to implant has softened the polymer matrix and pulled the surface
outward into an irregular wave pattern and has also displaced carbon fibers; (b) Bone has
osseointegrated with PMC inside an implant surface defect by degrading and replacing the
polymer matrix with osteoid that has substantially surrounded individual carbon fibers
approximately 50 μm into the implant.
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NIH-PA Author Manuscript NIH-PA Author Manuscript NIH-PA Author Manuscript
NIH-PA Author Manuscript NIH-PA Author Manuscript NIH-PA Author Manuscript
Petersen Page 27
Table 1
Biomaterial Properties
Material
Density (g/cm
3
)
Resistivity
a
(Ω m)
Tensile Strength (MPa) Yield Strength(MPa) Modulus(GPa)
Bone longitudinal (Ω m radial-longitudinal 100% wet) [1, 2] 1.8-2.1 45-150 90-149 114 15.2-18.6
Titanium grades 1-4 [1, 2, 14] 4.5-4.51
10
-7
240-550 170-485 104-110
Titanium-6-4aluminum vanadium alloy [1, 2, 14] 4.4-5.0
10
-8
860-1103 795-1034 116-120
Bisphenyl Unidirectional CF
b
[2, 14, 53, 54]
1.6 5 780-1850 140-325
Bisphenyl Unidirectional CF
b
4-pt. bend [2, 19, 29]
1.6 5 660-1800 64-255
Bisphenyl/CF
b
Exp.Uni-woven laminate 4-pt. bend [19, 29] 1.49 (±.01)
c
5
963 (±240)
c
774 (±176)
c
64 (±14.4)
c
Bisphenyl 3-D Woven E-Glass 3-pt. Bend X-Y planes [19]
576 (±129)
c
441 (±75)
c
26 (±18)
c
Unidirectional Photocure 3-pt Bend QF
b
[29] 1118.8 (±207.6)
c
76.6 (±13.3)
c
Polymer Acrylic Bone Cement (PMMA) 4pt Bend [14, 29] 1.17-1.20
>10
12
54.8 (±3.8)
c
43.2 (±3.6)
c
1.7 (±0.1)
c
a
Resistivity=1/Conductivity;
b
CF (Carbon Fiber), QF (Quartz Fiber);
c
Experimental standard deviations
Metals (Basel). Author manuscript; available in PMC 2015 January 27.
NIH-PA Author Manuscript NIH-PA Author Manuscript NIH-PA Author Manuscript
Petersen Page 28
Table 2
Resistivity
a
of Different Engineering and Biological Materials
Material Type Resistivity (Ωm)
Titanium Pure Conductor
4.2-5.2×10
-7
[14]
Titanium-6Al-4V Alloy Conductor
1.7×10
-8
[14]
Titanium Dioxide (rutile) Semiconductor 29-910 [65]
Bisphenol-Polymer/Carbon Fiber Composite Semiconductor 5 [19]
Bone Longitudinal Semiconductor 45-46 [2]
Bone Radial Semiconductor 150 [2]
Physiologic Saline Semiconductor 0.72 [2]
Silicon Pure Semiconductor 3000 [66]
Silicon Phosphorous Doped Semiconductor 20-80 [67]
Lipid Phosphate Headgroup/Water Interface Semiconductor 100 [64]
Carbon Fibers Conductor
9.5-18 × 10
-6
[14]
General Metals Conductors
~10
-6
-10
-9
[14]
Thermoset Bisphenyl Epoxy Polymer Insulator
10
10
-10
13
[14]
Acrylic Bone Cement Polymer Insulator
>10
12
[14]
Pure Quartz Fiber Insulator
10
20
[68]
a
Resistivity=1/Conductivity
Metals (Basel). Author manuscript; available in PMC 2015 January 27.
... et al. 2008). Modifications of metal surfaces have been employed to controlling tissue-titanium interactions; reduce the time for bone fixation and to prevent releasing of undesirable ions from the alloy (Richard C, 2015). These problems can be prevented by using coating layers from bioinert materials as bioceramic HAp. ...
... If large capillary distances or a local inflammatory hostile environment reduces oxygen availability then the implant passivation layer will fail to reform, leading to titanium corrosion. 95,102 Berbel et al 103 recently specifically assessed the effect of reduced oxygen availability for titanium repassivation in an electrochemical model of peri-implantitis and determined that reduced oxygen availability and inflammatory oxidizing conditions (eg, oxygen radicals) diminished titanium corrosion resistance. The hypothesis that titanium dissolution particles as products of corrosion, and not titanium wear particles due to implant insertion, are implicated in peri-implantitis is further supported by data from the orthopedic literature. ...
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