Polymerization kinetics stability, volumetric changes, apatite precipitation, strontium release and fatigue of novel bone composites for vertebroplasty
Polymerization kinetics stability, volumetric changes, apatite precipitation, strontium release and fatigue of novel bone composites for vertebroplasty
Piyaphong PanpisutID 0 1 2 3
Muhammad Adnan Khan 0 1 2 3
Kirsty Main 0 1 2 3
Mayda Arshad 0 1 2 3
Wendy Xia 0 1 2 3
Haralampos Petridis 1 2 3
Anne Margaret YoungID 0 1 2 3
0 Department of Biomaterials and Tissue Engineering, UCL Eastman Dental Institute , London , United Kingdom , 2 Department of Restorative Dentistry, Unit of Prosthodontics, UCL Eastman Dental Institute , London , United Kingdom
1 Editor: Tushar Jana, University of Hyderabad , INDIA
2 a Current address: Faculty of Dentistry, Thammasat University, Pathumthani, Thailand ?b Current address: Science of Dental Material, Khyber College of Dentistry, Khyber Teaching Hospital (MTI) , Peshawar , Pakistan
3 Royal Thai Government, Ministry of Sciences and Technology (
Data Availability Statement: All relevant data are
within the manuscript and its Supporting
The aim was to determine effects of diluent monomer and monocalcium phosphate
monohydrate (MCPM) on polymerization kinetics and volumetric stability, apatite precipitation,
strontium release and fatigue of novel dual-paste composites for vertebroplasty.
Materials and methods
Polypropylene (PPGDMA) or triethylene (TEGDMA) glycol dimethacrylates (25 wt%)
diluents were combined with urethane dimethacrylate (70 wt%) and hydroxyethyl methacrylate
(5 wt%). 70 wt% filler containing glass particles, glass fibers (20 wt%) and polylysine (5 wt
%) was added. Benzoyl peroxide and MCPM (10 or 20 wt%) or N-tolyglycine glycidyl
methacrylate and tristrontium phosphate (15 wt%) were included to give initiator or activator
pastes. Commercial PMMA (Simplex) and bone composite (Cortoss) were used for
comparison. ATR-FTIR was used to determine thermal activated polymerization kinetics of initiator
pastes at 50?80?C. Paste stability, following storage at 4?37?C, was assessed visually or
through mixed paste polymerization kinetics at 25?C. Polymerization shrinkage and heat
generation were calculated from final monomer conversions. Subsequent expansion and
surface apatite precipitation in simulated body fluid (SBF) were assessed gravimetrically
and via SEM. Strontium release into water was assessed using ICP-MS. Biaxial flexural
strength (BFS) and fatigue properties were determined at 37?C after 4 weeks in SBF.
Polymerization profiles all exhibited an inhibition time before polymerization as predicted by
free radical polymerization mechanisms. Initiator paste inhibition times and maximum
National Institute for Health Research (NIHR)
uk); The Engineering and Physical Sciences
Research Council (EPSRC: EP/I022341/1) (https://
epsrc.ukri.org); Wellcome Trust (ISSF/FHCE/0079)
(https://wellcome.ac.uk). This paper presents
independent research in part funded by the
National Institute for Health Research (NIHR)
under its Invention for Innovation (i4i) Programme
(Grant Reference Number II-LB-0214-20002). The
views expressed are those of the author(s) and not
necessarily those of the NHS, the NIHR or the
Department of Health. The funders had no role in
study design, data collection and analysis, decision
to publish, or preparation of the manuscript.
Competing interests: The work is covered by 2
licensed patents of Anne Young which may be
considered a conflict of interest as in the future she
may receive royalties when a commercial product
is produced. This does not alter our adherence to
PLoS ONE policies on sharing data and materials.
1. Formulations and materials with cationic
polymers (PCT/GB2014/052349, EP3027164A1,
US20160184190); 2. Formulations and composites
with reactive fillers (PCT/GB2007/003662,
EP2066703A1, EP2066703B1, US20100069469,
reaction rates were described well by Arrhenius plots. Plot extrapolation, however,
underestimated lower temperature paste stability. Replacement of TEGDMA by PPGDMA,
enhanced paste stability, final monomer conversion, water-sorption induced expansion and
strontium release but reduced polymerization shrinkage and heat generation. Increasing
MCPM level enhanced volume expansion, surface apatite precipitation and strontium
release. Although the experimental composite flexural strengths were lower compared to
those of commercially available Simplex, the extrapolated low load fatigue lives of all
materials were comparable.
Increased inhibition times at high temperature give longer predicted shelf-life whilst stability
of mixed paste inhibition times is important for consistent clinical application. Increased
volumetric stability, strontium release and apatite formation should encourage bone integration.
Replacing TEGDMA by PPGDMA and increasing MCPM could therefore increase suitability
of the above novel bone composites for vertebroplasty. Long fatigue lives of the composites
may also ensure long-term durability of the materials.
Osteoporotic fracture of the spine (osteoporotic vertebral fracture; OVF) causes severe pain,
height loss, limited mobility, kyphosis, and reduced pulmonary function [
treatments such as analgesics and rehabilitation are commonly used but often fail to relieve
severe pain in some patients [
]. Hence, surgical managements that relieve severe pain
rapidly such as vertebroplasty (VP) and balloon kyphoplasty (KP) are indicated. These procedures
involve injection of a bone cement to stabilize fractures. Common complications of these
treatments are cement leakage (up to 77%) leading to neurological deficits , adjacent vertebral
fractures (12?15%) [
], and post-operative infection which can be a rare but serious
Polymethyl methacrylate (PMMA) cement is the most commonly used bone cement for VP
and KP. Limitations of this cement include poor controlled setting and viscosity that may
increase the risk of cement leakage [
]. Further concerns are high polymerization shrinkage,
heat generation, and risk of toxic unreacted monomers release [
]. These shortcomings may
cause gap formation and local inflammation leading to fibrous encapsulation [
] reducing the
integrity of the bone-cement interface. Furthermore, conventional PMMA cements also lack
the ability to promote bone formation.
Two-paste injectable bone composites have been developed to address some limitations of
the PMMA cements but various shortcomings remain. For example, the primary base
monomer used has been bisphenol A-glycidyl methacrylate (Bis-GMA). This monomer is known to
limit final monomer conversion of composites due to its limited mobility [
the commercial composites contain TEGDMA as a diluent monomer, which is known to
increase shrinkage and heat generation of dental composites due to its high density of
methacrylate groups [
]. Furthermore, the composites contain the tertiary amine DMPT
(N,Ndimethyl-p-toluidine), which is highly toxic to human cells [
]. Recently developed
lightactivated urethane dimethacrylate (UDMA)-based dental composites exhibited higher
monomer conversion than Bis-GMA based commercial composites [
]. The same study also
2 / 25
demonstrated that replacing TEGDMA by polypropylene glycol dimethacrylate (PPGDMA)
increased monomer conversion and cytocompatibility of dental composites while
polymerization shrinkage was reduced.
Following mixing, chemically-activated bone composites should ideally cure rapidly after a
well-defined inhibition time that provides sufficient working time for injection. A potential
problem, however, is unmixed paste instability due to thermal initiated polymerization arising
upon storage [
]. Manufacturers usually recommend chemically-activated paste storage
below 4?C [
]. This ideal storage condition, however, may be difficult to achieve in some
circumstances. For example, medical products shipping to tropical regions can expose materials
to fluctuating temperatures between?4 to 42?C and 10 to 40?C during air and marine
transport respectively [
]. Stability may be estimated from inhibition times at elevated
temperatures for unmixed pastes. Temperature dependence of these times is expected to be governed
by Arrhenius type equations, be directly proportional to concentration of initiator (benzoyl
peroxide, BP) and inversely proportional to inhibitor levels added to stabilise different
The addition of monocalcium phosphate (MCPM) and tri calcium/strontium phosphates
(TCP/TSrP) into dental composites has been shown to promote hygroscopic expansion which
could potentially balance polymerization shrinkage and relieve residual shrinkage stress [
]. The addition of these reactive phosphates also enabled surface apatite formation which is
known to correlate with in vivo bone bonding [
]. Apatite formation can also be enhanced
by the addition of polylysine (PLS) [
]. Furthermore, strontium can promote osteoblast
proliferation and maturation whilst inhibiting osteoclast activities [
]. It will also enhance
], which may facilitate the surgical procedure and enable follow up with
Injected bone composites should be able to withstand the fluctuating and repetitive loads
during physical activities [
]. This may then help to prevent mechanical failure due to crack
propagation induced by repetitive subcritical loads (fatigue failure). A recent study [
indicated that high strength values of composites displayed under static loading were not directly
related to fatigue performance. Fatigue of various materials was previously assessed by
generating stress versus number of cycles curve (S-N curve) during simulated fatigue [
]. At a
given applied stress, the steep gradient of S-N plot was associated with a significant reduction
in failure cycles . Therefore, a low gradient rather than high gradient was preferable in
terms of fatigue performance [
The aim of this study was to compare TEGDMA/UDMA versus PPGDMA/UDMA-based
bone composites with added Ca/Sr phosphates (MCPM and TSrP) and polylysine (PLS).
Initiator paste stability, kinetics of polymerization, final monomer conversions, polymerization
shrinkage and heat generation, water sorption induced mass and volume changes, surface
apatite formation, strontium release, and biaxial flexural strength/fatigue were assessed. The effect
of MCPM levels (5 wt% versus 10 wt%) and type of diluent monomers (TEGDMA versus
PPGDMA) were examined.
Materials and methods
Material paste preparation
Experimental bone composites were prepared using a powder to liquid mass ratio of 70: 30.
The liquid phase (Table 1) contained urethane dimethacrylate (UDMA) (MW 479 g/mol,
DMG, Hamburg, Germany), polypropylene glycol dimethacrylate (PPGDMA) (MW 600 g/
mol, Polyscience, PA, USA) or triethylene glycol dimethacrylate (TEGDMA) (MW 286 g/mol,
DMG, Hamburg, Germany), and hydroxyethyl methacrylate (HEMA, MW 130 g/mol) (DMG,
3 / 25
Hamburg, Germany). To this was added either benzoyl peroxide (BP) (MW 242 g/mol
Polyscience, PA, USA) for the initiator liquid or N-tolyglycine glycidyl methacrylate (NTGGMA)
(MW 329 g/mol, Esschem, Seaham, UK) for the activator liquid.
Powder phase (Table 2) contained glass filler (particle diameter of 0.7 ?m, DMG, Hamburg,
Germany), glass fiber (30 ?m in diameter and 150 ?m in length, Mo-Sci, PA, USA),
monocalcium phosphate monohydrate (MCPM) particle diameter of 53 ?m, Himed, NY, USA),
tristrontium phosphate (TSrP) (particle diameter of 10 ?m, Sigma Aldrich, Gillingham, UK), and
polylysine (PLS) (particle diameter of 20?40 ?m, Handary, Brussel, Belgium).
Composite pastes were prepared at 23?C. Powders and monomers were weighed using a
four-figure balance (OHAUS PA214, Pine Brook, USA). The powder phase was mixed with
the liquid phase containing either initiator or activator using a planetary mixer (SpeedMixer,
DAC 150.1 FVZ, Hauschild Engineering, Germany) at 2000 rpm for 2 min. The initiator and
activator pastes were then poured into a double-barrel syringe (MIXPAC, SULZER,
Switzerland) over a vibrator to reduce air entrapment. The syringe was left in an upright position for
24 hr at 23?C to allow the release of air bubbles. For stability studies, pastes were then stored at
4, 23, and 37?C for 1, 3, 6, 9, and 12 months to give ?aged? pastes. Mixed experimental initiator
and activator pastes were obtained using an automatic mixing tip attached to the double-barrel
syringe and a mixing gun (MIXPAC Dispenser, SULZER, Switzerland). Commercial PMMA
cement (Simplex) and bone composite (Cortoss) (Table 3), mixed as per manufacturer?s
instructions within their use by date, were used as comparisons.
4 / 25
FTIR studies of composite pastes
Monomer conversion profiles. Monomer conversion profiles of pastes (n = 3) were
obtained using FTIR (Perkin-Elmer Series 2000, Beaconsfield, UK) with a temperature
controlled ATR attachment (3000 Series RS232, Specac Ltd., UK). Initiator pastes or mixed
experimental bone composites and commercial materials were placed in a metal circlip (1 mm depth
and 10 mm diameter) on the ATR diamond and covered with an acetate sheet. FTIR spectra
between 700?4000 cm-1 of the bottom surfaces of the specimens were recorded every 4 s at a
resolution of 4 cm-1. For unmixed pastes, spectra were recorded for up to 10 hours at 50, 60, 70
or 80 ? 1?C. With mixed pastes, spectra were obtained for 40 min at 25 ? 1?C.
Monomer conversion, Dc, and rate of polymerization, Rp, versus time were obtained from
FTIR spectra using Eqs 1 and 2 [
Where ?B0 and ?Bt were the absorbance of the C-O peak (1320 cm-1) above background level
at 1335 cm-1 initially and after time t and dDc/dt was the gradient of conversion versus time.
Furthermore, final monomer conversion, Dc,max, was calculated by linear extrapolation of
conversion versus inverse time to zero.
An example of a monomer conversion and reaction rate profile are shown in Fig 1A and
1B. These demonstrate a delay time (inhibition time) before rapid rise in monomer conversion
(snap set). Maximum rate is observed between 10?40% conversion with the reaction slowing
The standard mechanism of free radical polymerization of dimethacrylate monomers
includes, initiation, inhibition, propagation, crosslinking and termination steps. Using these
mechanisms, with the stationary state assumption that the concentration of free radicals is
Fig 1. Example profiles of A) polymerization and B) rate of polymerization of mixed M10PPG. All mixed and unmixed pastes exhibited
similar features for both profiles.
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constant, gives the inhibition time as [
[X] is the initial concentration of inhibitor and Ri the rate of initiation. Furthermore, rate of
polymerization (Rp) can be described using the following equation[
Rp ? kp?M Rkti
[M] is the monomer concentration and k fi
pffpfafnidifkt rate constants for propagation and
termination steps. Combining Eqs 3 and 4, Rp ti is therefore expected to be independent of rate of
initiation and given by the following equation.
Rp ti ? kp?M
In the following, inhibition times were calculated by linear extrapolation of data between
10% and 40% monomer conversion back to 0% conversion. The gradient in this range was
used to obtain the maximum rate of polymerization Rp,max and Rp;max ti.
Thermally activated polymerization of unmixed initiator paste, activation energies and
predicted shelf life. Pilot studies revealed that initiator pastes were more susceptible to heat
than activator pastes and that modifying MCPM level had relatively minimal effect. Hence,
initiator pastes of M10PPG and M10TEG were chosen to assess polymerization kinetics and
thermally activated polymerization of the experimental bone composites. FTIR spectra of freshly
mixed M10PPG and M10TEG initiator pastes (n = 1) were used to obtain their inhibition times,
rates of polymerization and final conversions at temperatures of 50, 60, 70, and 80?C.
Inverse inhibition times and reaction rates are proportional to rate constants, k, whose
temperature dependence, are generally described by Arrhenius type equations [
ln k ? ln A
T is temperature in Kelvin and R the gas constant. A is a pre-exponential factor that is related
to the frequency of molecular collisions between reacting species and Ea, the activation energy
required for them to react.
Combining Eqs 3, 4 and 6, ln(1/ti) or lnRp,max versus 1/T are expected to be linear if Ea is
temperature independent. In the following, these were plotted, and used to obtain activation
energies for the initiation step and monomer conversion respectively. These plots were also
extrapolated to estimate times of inhibition and 50% monomer conversion at 4, 23, and 37?C.
These times provided estimates of when pastes might be expected to start polymerizing and
Visually observed solidification of unmixed pastes and stability of mixed paste
polymerization kinetics (observed shelf life). To visually assess paste hardening with long-term
storage, double-barrel syringes containing unmixed M10PPG and M10TEG initiator and
activator pastes were stored at controlled temperatures of 4, 23, and 37?C. At 1 day, 1, 3, 6, 9 and
12 months, small portions were extruded to check for solidification. To assess stability of
mixed paste polymerization kinetics at these times, for the 4?C stored samples a portion of the
composite was mixed and polymerization kinetics at 25?C determined by FTIR-ATR (n = 3).
Polymerization kinetics, shrinkage and heat generation of freshly prepared and mixed
pastes. To compare polymerization kinetics of different mixed composites, inhibition times,
maximum reaction rates and final conversions of freshly prepared and immediately mixed
experimental materials were compared with those for commercial PMMA (Simplex) and bone
composite (Cortoss) cements. Additionally, for the experimental materials, polymerization
volume shrinkage (?) (%) and heat generation ( )(kJ/cc) were calculated using the following
??%? ? e Mf Dc;max?
? ? ? Mf ?D1c0;m0ax??
where Mf, monomer mass fraction; Dc, monomer conversion (%); ?, composite density (g/
cm3); ni; the number of C = C bonds per molecule; Wi, molecular weight (g/mol) of each
monomer; xi, mass fraction of each monomer in the liquid. These assume one mole of
polymerizing C = C groups gives volumetric shrinkage of 23 cm3/mol (e) and generates 57 kJ of
heat (?) .
Properties of set discs prepared from fresh pastes
Polymerized disc preparation
To produce disc samples, freshly prepared and then mixed pastes were injected into metal
circlips (1 mm in thickness and 10 mm in diameter). The samples were covered with an acetate
sheet on top and bottom surfaces. The samples were left for 24 hr at 23?C to allow completion
of polymerization. After removal from circlips, any excess was carefully trimmed. The set
samples were subsequently immersed in a tube containing 10 mL of simulated body fluid (SBF)
prepared according to BS ISO 23317:2012  or deionized water at 37?C until the required
Mass and volume changes
Mass and volume changes of set composite discs after immersion in SBF at 37?C for 0, 1, 6, 24,
48 hr and 1, 2, 3, 4, 5, 6 weeks were measured using a four-figure balance with a density kit
(Mettler Toledo, Royston, UK). The percentage mass and volume change, ?M and ?V, were
determined using the following equations .
where M0 and V0 is initial mass and volume, whilst Mt and Vt are mass and volume at time t
Surface apatite formation
To assess ability of materials to promote surface apatite formation, disc specimens were
immersed in SBF and incubated at 37?C for 1 week (n = 1). They were subsequently removed
and sputtered with gold-palladium using a coating machine (Polaron E5000, East Sussex, UK)
for 90 s at 20 mA. The specimens were examined under SEM (Phillip XL-30, Eindhoven, The
Netherlands) operating with primary beam energy of 5 kV and a current of approximately 200
Strontium (Sr2+) release
Sr2+ release was measured from experimental bone composites discs (n = 3) immersed in 10
mL of deionized water. The specimens were incubated at 37?C for 4 weeks. The storage
solution was collected and replaced with a fresh solution at 24 hr, 1, 2, 3, and 4 weeks. The collected
solution was mixed with 2 vol% nitric acid (1:1 volume ratio). Calibration standards
containing Sr2+ of 1 ppb, 2.5 ppb, 10 ppb, 25 ppb, 100 ppb, 250 ppb, and 1 ppm were prepared using
the ICP-multi element standard solution XVI (Certipur Reference Materials, Merck KGaA,
Germany). The cumulative Sr2+ release was calculated using the following equation.
%Sr release ?
where wSr is the initial amount of Sr2+ in the sample (g), St is the amount of Sr2+ released into
storage solution (g) collected at time t (hr).
Biaxial flexural strength and fatigue life
To assess biaxial flexural strength (BFS) and fatigue performance of the materials, disc
specimens were immersed in SBF and incubated at 37?C for 4 weeks (n = 25). Prior to fatigue
testing, BFS of the composites was assessed using a ?ball on ring? jig with a servo hydraulic testing
frame (Zwick HC10, Zwick Testing Machine Ltd., Herefordshire, UK) equipped with a 1 kN
load cell (n = 5) [
]. The specimens? thickness was recorded using a digital vernier calliper.
The sample was placed on a ring support (8 mm in diameter). The load was applied using a 4
mm diameter spherical ball indenter at 1 mm.min-1 crosshead speed. The failure stress was
recorded in newton (N) and the BFS (S; Pa) was calculated using the following equation [
F n h i o
S ? 2 ?1 ? ?? 0:485ln ? 0:52 ? 0:48
Where F is the load at failure (N), d is the specimen?s thickness (m), r is the radius of circular
support (m), and v is Poison?s ratio (0.3).
For assessing fatigue performance, a sinusoidal load of 5 Hz [
] was applied to specimens
using 80%, 70%, 60%, and 50% of mean BFS (n = 20, 5 specimens per each level of stress). The
tests were continued until fracture occurred or the requisite number of load cycles (100,000
cycles) had been applied. BFS was plotted against cycles of failure to generate classical
stressnumber of cycle curve (S/N curve). Failure cycle from 1st to 5th samples were plotted against
BFS which therefore gave five S/N curves. Mean of gradient from the plots was obtained
(n = 5) and used to compare fatigue performance [
]. Furthermore, number of failure cycles
(fatigue life) at stress level of 10 MPa was obtained by extrapolating the regression line. This
value represents fatigue life of materials upon applying low stress that may be generated during
normal movements such as flexion, lateral bending, or walking .
All values and errors reported throughout this study were mean (? standard deviation; SD).
SPSS Statistics software (version 24 for Windows, IBM, USA) was used for statistical analysis.
Homogeneity of variance was assessed using Levene?s test. When variances were equal, data
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were analyzed using one-way analysis of variance (ANOVA) followed by post-hoc Tukey?s test
for multiple comparisons. Alternatively, the Kruskal-Wallis test, followed by multiple
comparison using Dunnett?s T3 tests, was used if the variances were not equal . Correlation
between inhibition time and final monomer conversion with storage time was tested using
Pearson?s correlation. The significance value was set at p = 0.05. Line fitting for regression
analysis was undertaken using the Linest function in Microsoft Office Excel 2016 for
Factorial analysis was used to assess the effect of MCPM level and diluent monomers on
properties of composites from freshly prepared pastes. A full factorial equation for two
variables each at high (10 wt% MCPM, PPGDMA) and low levels (5 wt% MCPM, TEGDMA) can
be fitted using the following equation [
lnP ?< lnP > a 1
Where a1 and a2 were the effect of each variable on the property P of the composites, < lnP >
is the average value of lnP. The a1,2 term is an interaction effect. The percentage effect of each
variable, Q, can be calculated using the following equation.
Q?%? ? 100 1
GH and G0 are the geometric average property for the samples with the variable at its high
versus low value respectively. The effect of variable change was considered significant if the
magnitude of ai was greater than both its calculated 95% CI and interaction terms.
FTIR studies of composite pastes
Thermally activated polymerization of unmixed initiator paste, activation energies and
predicted shelf life. An example of M10PPG initiator paste conversion versus time and
temperature is shown in Fig 2A. Upon raising temperature from 50 to 80?C, averageptiffdeffcrieifased
from 27,000 to 350 s, whilst Rp.max increased from 0.010 to 0.23%/s giving Rp;max ti of 1.9 to
4.3%s-0.5. Profiles for M10TpEGffefxfhiiifbited shorter inhibition times of 1,600 to 70 s, Rp.max of
0.011 to 0.44%/s and Rp;max ti of 0.4 to 3.6%s-0.5.
ln(1/ti) and ln(Rp,max) versus inverse temperature (Fig 2B) gave linear plots (R2 > 0.99).
M10PPG initiation and polymerization activation energies calculated from these were 137 and
95 kJ/mol whilst lnA terms were 41 and 31 respectively. For M10TEG, activation energies were
100 and 120 kJ/mol and lnA terms were 30 and 40 respectively. Extrapolation, gave M10PPG, ti
of 3 days, 1 month, and 51 months at 37, 23, and 4?C respectively. Times for 50% conversion
were comparable. Conversely, M10TEG ti was 3 hours, 18 hours and 12 days, whilst times for
50% conversion, were 19 hours, 7 days and 7 months at 37, 23, and 4?C respectively.
At 50% conversion, the reaction rates began to slow and tended to final conversions that
declined linearly versus 1/T (Fig 2B). At 50?C, reaction level following the 10 hours of
observation was too low with M10PPG to enable determination of final conversion. Between 80 and
60?C, however, final monomer conversion declined from 94 to 81% for M10PPG and from 90
to 73% for M10TEG.
Visual solidification of unmixed pastes and stability of mixed paste polymerization
kinetics (observed shelf life). Visual inspection indicated that, at 37?C, both M10PPG and
M10TEG initiator pastes solidified in the syringes between 1 day and 1 month. At 25?C,
M10PPG polymerized between 3 and 9 months whilst M10TEG initiator pastes solidified
9 / 25
Fig 2. A) Example (n = 1) polymerization profiles of unmixed initiator paste of M10PPG at different temperatures.
The paste contains 3 wt% BP and 20 wt% MCPM with no activator nor TSrP. B) Average (n = 1) inhibition time (ti;
circles), Rp,max (diamond), and final monomer conversion (triangle) of M10TEG and M10PPG initiator pastes plotted
as Ln(1/ti) or Ln(Rp,max) versus inverse of temperature (Arrhenius plots). Reaction rates at a given temperature were
largely similar for M10TEG compared with M10PPG although delay times and final conversions were significantly and
slightly reduced respectively.
between 1 and 3 months. All initiator pastes stored at 4?C, however, remained fluid even after
12 months storage.
Fig 3 shows inhibition time and monomer conversion of mixed M10PPG and M10TEG
pastes that had been stored at 4?C unmixed for up to 12 months. With M10PPG, inhibition
time (74 s), polymerization rate (0.43%/s), Rp;max ti (3.7%s-0.5i) and final conversion (79%)
exhibited only minor change with pre-aging of pastes. With M10TEG, Rp;max ti and final
conversion also remained constant at 2.1%s-0.5 and 70% respectively. M10TEG inhibition time,
however, showed a significant increase from 53 s at 24 hr to 104 s at 12 months of pre-aging
(R2 = 0.77, p = 0.02).
Polymerization kinetics of freshly prepared and mixed materials. Increasing MCPM
level showed no significant effect on inhibition time and rate of polymerization. The shortest
10 / 25
Fig 3. Inhibition time (diamond) and final monomer conversion (triangle) of the mixed experimental bone composites after storage
unmixed at 4?C for up to 12 months. Error bars are SD (n = 3 at each time point).
and longest inhibition times were observed with M10TEG (24 ? 4 s) and Simplex (496 ? 17 s)
respectively (Fig 4A). Inhibition times of all experimental bone composites (24?96 s) were
shorter than that of Simplex and Cortoss (169 ? 23 s). Average inhibition time of PPGDMA
based composites (85 s) was longer than that of TEGDMA based composites (24 s). Factorial
analysis indicated that replacing TEGDMA by PPGDMA increased inhibition time by ~ 250%
whilst the effect of increasing MCPM level was negligible. Additionally, experimental
ites exhibited comparable Rp,max to commercial materials (Fig 4B). Average Rp;max ti for
PPGDMA, TEGDMA, Cortoss and Simplex were 4.7, 2.5, 5.4 and 10.9%s-0.5 respectively.
Final monomer conversions (Dc,max) of experimental bone composites were higher than
that of Cortoss (64 ? 1%) (Fig 4C). M5PPG (82 ? 1%) showed significantly higher final
monomer conversion than Simplex (78 ? 1%) (p < 0.05). Averaged final monomer conversion of
PPGDMA-based bone composites (80%) was greater than that of TEGDMA-based composites
(76%). Factorial analysis indicated that replacing TEGDMA by PPGDMA increased monomer
conversion on average by 5%. Additionally, monomer conversion of the composites was
increased by 5% upon decreasing MCPM level from 10 to 5 wt%.
Average calculated polymerization shrinkage and heat generation of TEGDMA-based
experimental bone composites (5 vol% and 0.13 kJ/cc) were slightly higher than those of
PPGDMA-based composites (4 vol% and 0.10 kJ/cc) (Fig 4D). Factorial analysis indicated that
the calculated shrinkage and heat generation were increased by 22% upon replacing PPGDMA
11 / 25
Fig 4. A) inhibition time, B) maximum rate of polymerization, C) final monomer conversion for experimental and
commercial products, and D) calculated polymerization shrinkage and heat generation for experimental composites
tested at 25?C. Lines indicate no significant difference (p > 0.05). Error bars are SD (n = 3).
Mass and volume changes
Initial mass and volume change of materials increased linearly with square root of time
consistent with diffusion-controlled water sorption. Simplex and Cortoss equilibrium values of mass
change were 1.6 ? 0.1 wt% and 3.0 ? 0.1 wt% (Fig 5A and 5B). Mass changes at later times of
3.4 ? 0.1 wt% (M5TEG), 4.0 ? 0.1 wt% (M5PPG), 4.3 ? 0.2 wt% (M10TEG), and 5.8 ? 0.0 wt%
(M10PPG) were obtained for the experimental materials. Replacing TEGDMA by PPGDMA
and increasing MCPM level increased mass changes at late time by 34 ? 1% and 27 ? 3%
Volume change of Simplex and Cortoss reached maximum values at 1 week of 1.0 (? 0.1)
and 2.8 (? 0.2) vol% (Fig 5C and 5D). Later time maximum values were 1.5 (? 0.1), 3.0 (? 0.3),
4.3 (? 0.7), 6.1 (? 0.3) vol% for M5TEG, M5PPG, M10TEG, and M10PPG respectively. Factorial
analysis indicated that replacing TEGDMA by PPGDMA and rising MCPM level enhanced
volume change at late time by 45 ? 15% and 109 ? 8% respectively.
Surface apatite formation
At 1 week, no precipitates were observed on surfaces of M5TEG, Simplex, and Cortoss (Fig
6). Patchy crystals consistent with brushite were observed on some areas of M5PPG.
Conversely, thin patchy surface apatite (~ 1 ?m) layers partially covered surfaces of M10TEG and
12 / 25
Fig 5. Mass and volume changes versus square root of time (hr) of all materials immersed in SBF for up to 6 weeks. Error bars are SD
(n = 3).
The cumulative release of Sr2+ increased linearly with time (hr) (Fig 7). Highest and lowest
rate of Sr2+ release was 0.0015%.hr-1 and 0.0006%.hr-1 observed with M10PPG and M5TEG
respectively. M10PPG exhibited the highest accumulative Sr2+ release at 4 weeks (1.12 ?
0.02%). Factorial analysis indicated that cumulative release of Sr2+ at 4 weeks was increased by
127 ? 14% upon increasing MCPM level. Additionally, the release was increased by 111 ? 42%
upon replacing TEGDMA by PPGDMA.
Biaxial flexural strength and fatigue
The highest and lowest BFS of materials tested in SBF at controlled temperature of 37?C was
obtained from Simplex (137 ? 4 MPa) and M10PPG (54 ? 2 MPa) respectively (Fig 8A).
M5PPG had a comparable BFS (63 ? 3 MPa) to M5TEG (67 ? 4 MPa). The BFS of M5TEG was
significantly higher than that of M10PPG (54 ? 2 MPa), M10TEG (57 ? 2 MPa), and Cortoss
13 / 25
Fig 6. Representative SEM images for each material after immersion in SBF for up to 7 days.
(58 ? 2 MPa). Factorial analysis showed that BFS was increased by 18 ? 5% upon decreasing
MCPM level, whilst the effect of diluent monomers was negligible.
BFS versus logarithm of failure cycle number (S/N curve) gave straight line plots with
negative gradients (Fig 8B). The most negative gradient was observed with Simplex (-17.7 ? 2.6
MPa/log cycle) (Fig 8C). The gradients for M5PPG (-6.7 ? 0.5 MPa/log cycle), M10PPG
(-5.9 ? 1.5 MPa/log cycle), M5TEG (-5.9 ? 1.4 MPa/log cycle), M10TEG (-5.6 ? 0.8 MPa/log
cycle) and Cortoss (-6.0 ? 0.3 MPa/log cycle) were comparable. Factorial analysis indicated
that the effect of MCPM level and diluent monomers were negligible.
Extrapolated failure cycle values (fatigue life) at 10 MPa for the experimental composites
(7.5?8.2 log cycle) were not significantly different from that for the commercial materials (7.8?
7.9 log cycle) (Fig 8D). Additionally, factorial analysis showed that MCPM level and diluent
monomer had no significant effect on the fatigue life.
14 / 25
Fig 7. Cumulative Sr2+ release versus hr from bone composites immersed in deionized water up to 4 weeks. Error bars are SD (n = 3).
This study produced bone composites that are two-paste, chemical-cured versions of
previously developed Ca/Sr and PLS-containing, single paste light-cured dental composites [
The main change involved light activated initiator (camphorquinone) replacement with a
chemical activated initiator (BP). This enables the composite to cure chemically after mixing
with a separate amine activator-containing paste instead of following light activation.
Additionally, however, the tertiary amine activator N,N-dimethyl-p-toluidine (DMPT) was
replaced by polymerizable NTGGMA to reduce the risk of toxic amine activator leaching.
Furthermore, PLR was reduced from 4:1 to 2.3:1 to enable easy mixing of the initiator and
activator-containing pastes through a fine mixing tip and enhanced flow within the vertebra. The
effect of MCPM level and diluent monomers on various chemical and mechanical properties
FTIR studies of composite pastes
The shelf life of chemically-activated bone cement is affected by storage temperature,
monomer type, inhibitor, initiator and activator levels [
]. Unmixed composite paste stability is
crucial to avoid premature or thermal initiated polymerization during storage or shipment.
15 / 25
Fig 8. A) BFS tested in SBF at 37?C, B) example plots of BFS versus log (cycle) (n = 20), C) gradients of S/N plots in
positive values for clarity purpose, and D) extrapolated fatigue life at BFS of 10 MPa. Lines indicate no significant
difference (p > 0.05) and error bars are SD (n = 5).
Furthermore, polymerization kinetics following mixing must be stable and controllable to
enable effective setting under clinical conditions.
In this study, a temperature-controlled FTIR-ATR system was employed to monitor
polymerization kinetics. As paste at room temperature was placed on the hotter ATR plate and
reaction kinetics are highly sensitive to temperature a potential error arises due to the time
taken for the paste to reach the ATR temperature. This is reduced through use of thin samples.
For the elevated temperature studies, this error was further minimized through ensuring
inhibition times were more than 60 s. As reactions were monitored for up to 10 hours, this enabled
a wide range of reaction temperatures and times. The reaction temperature for mixed pastes of
25?C mimics clinical conditions before injection into the vertebra and being just slightly above
room temperature minimizes temperature variability errors.
In order to understand and predict kinetics of mixed and unmixed composite paste
polymerization under different conditions, reaction mechanisms and theories were employed. The
mechanism for dimethacrylate reaction used in the derivation of Eqs 3 to 5 included initiation,
inhibition, propagation, crosslinking and bimolecular termination steps  which may be
I ! 2R
R ?X ! R ? X
R ?M ! M
Mn ?M ! Mn?1
R ?Mn ! Mn
Mn ?Mm ! Mn?m
16 / 25
A limitation of the theory is the possibility of inhibition via routes other than by the added
inhibitor. This could include free radical loss by oxygen inhibition or upon contact with
surfaces such as of the filler particles or the container that may contain inhibitors [39, 40]. The
termination step may also occur via routes other than through bimolecular collision of two
polymer free radicals . Changes in relative importance of different mechanisms with
reaction rate could cause errors in prediction of lower temperature stability. Lower temperature
reaction rates predicted from Arrhenius plots were therefore compared with the stability and
reaction kinetics of pastes that had been stored long-term.
Paste stability and inhibition times. The observation of a delay time prior to rapid
polymerization is expected from kinetic theories for both thermal and amine activated reactions
]. For initiator and mixed pastes, this will indicate their shelf-life and time available for
injection into the body (working time) respectively. According to Eq 3, the inhibition time is
proportional to the inhibitor concentration and inversely proportional to initiation rate. In
initiator pastes, initiation rate is proportional to the benzoyl peroxide concentration. For amine
activated reactions it is proportional to initiator and activator concentrations [
]. To increase
initiator-paste stability and mixed paste working time, the inhibitor can therefore be increased,
or the initiator and activator reduced.
Extrapolated Arrhenius plots predicted the greater low temperature stability of the
PPGDMA compared with the TEGDMA initiator paste. Calculated shelf-lives, however, were
~10 times lower than those observed through long-term paste storage. This may be due to the
alternative mechanisms of inhibition such as by oxygen or surface of the container when the
rate of polymerization is slow. Calculated initiator paste inhibition times at 23?C were ~104
greater than those for the mixed pastes. Replacing half of the initiator by activator through
paste mixing, therefore enabled rapid setting of the composite pastes.
The inhibitor in the supplied diluent TEGDMA monomer was 200 mM of MEHQ
(4-methoxyphenol), whilst that of PPGDMA monomer was a mixture of 100 mM of MEHQ and 100
mM of BHT (butylated hydroxytoluene). A previous study has demonstrated that the addition
of BHT enhanced the stabilisation effect of MEHQ  which might explain in part the
observed lower inhibition times and stability of the TEGDMA initiator pastes.
The lower pre-exponential term and activation energy for the initiation step predicts faster
free radical production with the TEGDMA initiator pastes at lower temperature but vice versa
at high temperature. A possible explanation is that the smaller size of TEGDMA molecules
reduces initial steric hindrance thereby lowering the activation energy for formation of free
radicals when compared with UDMA. Conversely the larger PPGDMA molecules are of
comparable size to the bulk UDMA possibly giving more comparable activation energies for
free radical formation. Higher concentrations of reacting molecules but slower monomer
radical formation in the PPGDMA pastes might then explain the differences in reaction kinetics.
Polymerization rates. According to Eq 4, the rate of polymerization following the
inhibition period is proportional to the rate of initiation. Preferably rate of polymerization following
injection of a mixed paste should be rapid to prevent leakage from the injection site or
subsequent release of monomers in the body. Rate of polymerization can potentially be raised by
increasing the initiator and activator concentrations. The inhibitor may then additionally need
to be raised to maintain the required working time and initiator paste stability. The rate of
polymerization of the mixed pastes was comparable with that of the initiator pastes at 80?C
and ~250 times that predicted for the unmixed initiator pastes at room temperature.
From Eq 5, Rp;max ti might be expected to be constant. Rp;max ti values for mixed pastes
were comparable with those for the initiator pastes at the highest temperatures. A possible
explanation for the decrease in Rp;max ti for the initiator paste with decreasing temperature,
17 / 25
however, could be alternative free radical inhibition and termination reactions when the
reaction is slow. These alternative radical removal reactions might also result in loss of the benzoyl
peroxide initiator upon storage. This could then explain the increase in inhibition time of
mixed pastes following long-term TEGDMA initiator paste storage.
The higher pre-exponential term for the polymerization propagation step observed with
the TEGDMA initiator pastes is to be expected if higher concentration of free radicals are
generated. The higher activation energy for the propagation step may be due to the TEGDMA
radicals requiring more energy than UDMA or PPGDMA radicals to react with UDMA
Maximum monomer conversions. Following 50% monomer conversion, the slowing of
the dimethacrylate reaction rates can be explained by the propagation reaction that generates
linear polymer chains, changing to a crosslinking process. The reaction will slow further when
the conversion is sufficient to convert the material from a crosslinked rubber into a solid glassy
polymer . At elevated temperatures, higher conversion is required for this glass transition
temperature to be reached.
Final conversions at room temperature for the PPGDMA and TEGDMA composites are
comparable with values obtained using the same monomers but light activated polymerization
]. Greater final conversion with the PPGDMA pastes could be a consequence of the longer
flexible polypropylene glycol chain lowering the glass transition temperatures. Additionally, if
the reaction is continuing at a fast rate when it solidifies, high concentrations of free radicals
and localized heating could enable higher conversion. With the mixed fast reacting pastes,
conversions at 25?C were comparable with those achieved at 60?C with the slower reacting
Polymerization of freshly prepared materials. Working time of PMMA bone cements
that require mixing powder with liquid and transfer to a syringe for vertebroplasty should be
approximately 6 to 10 min . Approximately 3 minutes is required for mixing. 4?8 mL of
PMMA cement is generally sufficient to stabilize a fractured vertebra . With an injection
rate of 0.15 mL/s , an injection time of 0.5?1 minutes is then required to deliver bone
cement through a cannula to an affected site. This must be undertaken before the paste
viscosity becomes too high for injection. This change in viscosity occurs due to swelling of the beads
in the monomer phase.
For a two-paste bone composite in double-barrel syringe, the mixing takes only a few
seconds. Additionally, no change in rheological properties occurs following mixing, lower
volumes are required to stabilise fractures, and less heat generation compared with PMMA
cement . These features are a distinct advance for the composites and enable significant
shortening of the required working time.
The inhibition times measured from FTIR-ATR following mixing of both the
powder-liquid PMMA cement (496 s) and two-paste Cortoss bone composite (169 s) are different to final
setting times cited in the literature  (378 s for Simplex and 345 s for Cortoss). This may be
a consequence of using a different method (surface indentation), volumes of material in the
test and batch number or time after production. The inhibition time of freshly prepared
TEGDMA based composite was too short (23 s) indicating that more inhibitor should be
included. According to Eq 3, the inhibitor concentration would need to be increased 7 folds in
order to bring the inhibition time up to that of Cortoss. This might additionally enhance the
initiator paste shelf-life. With the PPGDMA paste, a doubling in inhibitor should give a similar
inhibition time to that of Cortoss. From Eq 4, increased inhibitor should not affect the rates of
polymerization. The similarities in experimental and commercial material reaction rates
suggests this change would thus enable production of composites with ?snap set? following
sufficient working time for injection.
18 / 25
High final monomer conversion is required for good physical/mechanical composite
properties in addition to the low risk of toxic monomer leaching [50, 51]. The final monomer
conversion of Cortoss in the current study was lower than that previously obtained (80%) using
differential scanning calorimetry (DSC) . DSC, however, has given higher final conversion
compared to FTIR in other studies [
]. Lower monomer conversion of Cortoss compared
with experimental bone composites could be due to different primary base monomers.
Generally, high glass transition temperature (Tg) monomers give low final monomer conversion [
]. Primary base monomer of Cortoss is Bis-GMA (Tg = -7.7?C), whereas that of the
experimental bone composites is UDMA (Tg = -35.3?C) .
Monomer conversion of Simplex (77%) in the current study is in good agreement with that
obtained from published studies (70%) [
]. Simplex contains the monomethacrylate,
methyl methacrylate (MMA), which unlike dimethacrylates, gives only linear chains and no
crosslinking reaction. Consequently, complete polymerization of all methacrylate groups is
required to prevent monomer leaching. Conversely, with dimethacrylate-containing
composites, 50% conversion may be sufficient to bind all monomers within the resin matrix . Hence
70?80% observed conversion with the experimental bone composites is expected to reduce the
risk of unreacted monomer release and potentially lead to improved cytocompatibility [
Polymerization heat generation and shrinkage. The lower concentration of double
bonds per mole of PPGDMA contributed to the lower calculated polymerization shrinkage
and heat generation of PPGDMA-based composites compared to the TEGDMA-based
composites. The shrinkage of experimental bone composites in the current study was comparable
to that of Cortoss (5 vol%) [
] but lower than that of PMMA bone cement (6?7 vol%) [
Additionally, as heat generation is proportional to shrinkage and a lower volume of composites
is required to stabilize vertebral fractures , the composites should cause less thermal
damage than Simplex upon placement. These properties may help to minimize gap or fibrous
capsule formation and improve interfacial integrity at the bone-composite interface.
Mass and volume changes
Mass increase due to water sorption of Simplex reached equilibrium within 1 week which is in
accordance with a published study [
]. Cortoss exhibited greater mass increase compared to
Simplex due probably to the lower monomer conversion, higher flexibility of polymer
network, and hydrophilicity of bioactive glass contained in the composite [
The volume increase of Simplex in the current study (~ 2 vol %) was lower than the
polymerization shrinkage reported from a published study (6?7 vol%) [
]. This mismatch
between shrinkage and expansion may cause gap formation and induce fibrous encapsulation
at the bone-cement interfaces. This poor material-bone integration may impair load transfer
mechanisms leading to re-fracture or progression of cracks toward adjacent vertebra [
For experimental bone composites, their mass and volume changes were governed primarily
by MCPM level and type of diluent monomer. Raising MCPM level enhanced water uptake
leading to the increase of mass and volume as was previously observed with dental composites
]. Low crosslinking density due to the high molecular weight of PPGDMA could promote
water diffusion, thereby increasing the mass and volume changes of the PPGDMA-based
composites. For PPGDMA formulations, therefore, 5 to 10 wt% of MCPM was sufficient to enable
hygroscopic expansion comparable with the calculated polymerization shrinkage. TEGDMA
formulations, however, may requires greater than 5?10 wt% of MCPM to allow expansion to
compensate polymerization shrinkage. These expansions are expected to relieve shrinkage stress
and minimize gaps at composite-bone interface. This could potentially help to improve
interfacial integrity and load transfer and reduce recurrent fracturing of the treated vertebra.
19 / 25
Surface apatite formation
The apatite-forming ability in SBF has been adopted as a method for the determination of the
bone bonding potential in biomaterials prior to any animal testing which requires large
expenses and resources. It is proposed that the formation of surface apatite is associated with
the ability of materials to promote in vivo bone bonding [
]. Other studies with
MCPM-containing composites demonstrated that the level of apatite precipitation increased
proportionally to time [
]. Mineral release is also expected to promote mineralization of newly formed
When surface MCPM dissolves, it disproportionates into phosphoric acid and dicalcium
phosphate. Under acidic conditions, the later will precipitate as brushite. If the acid is
neutralized by buffering ions in the SBF, the brushite can transform into apatite [
MCPM level from 5 to 10 wt% encouraged this transformation and greater precipitation of
surface apatite after immersion in SBF for 1 week. Replacing TEGDMA by PPGDMA,
however, provided no obvious advantageous effect on surface apatite precipitation. In Cortoss,
bioactive glass was added in an attempt to provide mineralization and enhance bonding with
bone . Surface apatite was however not seen after SBF immersion for 1 week. This could be
due to the slower dissolution of its calcium phosphate containing glass (combeite) [
compared with MCPM.
Strontium ion release
Strontium release is of interest due to its potential beneficial effects for bone repair including
increase of osteoblast proliferation and reduction of osteoclastic activities [
22, 23, 68
observed linear release of Sr2+ from experimental bone composites suggests it is not
diffusioncontrolled. It is possible that the level of strontium release was dependent upon its release from
the surface following water enhancing polymer expansion. This would explain the increase of
strontium release observed upon using PPGDMA and rising MCPM level. The results gave the
effect of replacing TEGDMA by PPGDMA on Sr2+ release as less than the effect upon doubling
MCPM level. It is hypothesized that in addition to increased water sorption, MCPM may
produce phosphoric acid that reacts with the tristrontium phosphate to form distrontium
phosphate of higher aqueous solubility and thereby higher Sr ion release.
The release of Sr2+ would be enhanced upon increasing surface area. A previous study
showed that a bone composite provided greater interfacial stability at bone/material interface
than a PMMA cement [
]. This was attributed to possibly a faster bone response, improved
bone binding to mineral precipitation around the composite, and/or more effective
penetration of the composite into porous bone. The greater penetration could give a large surface and
encourage greater localized release of strontium to the surrounding osteoporotic vertebra.
This may potentially help to increase bone mass and improve mechanical properties of the
vertebra, thereby decreasing the risk of recurrent fractures. The release of strontium that increases
linearly with time may also enable constant drug release which may be considered beneficial.
Biaxial flexural strength (BFS) and fatigue
The mean BFS values obtained from Simplex and Cortoss were consistent with those reported
in a previous study . Mean BFS of experimental bone composites was also comparable to
that of Cortoss. Increasing hydrophilic contents and flexibility of polymer networks usually
reduces strength of composites. Results from the current study showed that increasing MCPM
level and replacing TEGDMA by more flexible PPGDMA had no significant effect on the
strength and fatigue of the composites. This might be due to the low level of MCPM used and
the enhanced monomer conversion from PPGDMA. Despite the fact that specimens were
20 / 25
aged for 4 weeks, the mean BFS values of all experimental bone composites were greater than
the 24 hr flexural strength of 50 MPa required by ISO 5833: Implants for surgery?Acrylic
resin cements [
The highest gradient of S-N curve was observed with Simplex. This could be due to the lack
of reinforcing glass fillers or glass fiber to retard crack propagation. Additionally, pores caused
by the poor integration between BaSO4 particles and polymer matrix could also act as crack
]. It is assumed that the lower gradient of S-N curve of Cortoss and experimental
bone composites compared with Simplex may result from the beneficial effects of absorbed
water that could improve fatigue resistance. The water can plasticize resin matrix and increase
polymer chain mobility which could enhance crack tip blunting [
]. For experimental bone
composites, releasing of active ingredients may leave voids behind but the contained fibers
could help to bridge the voids and slow down crack initiation [
In physiologic conditions, the injected bone cements are expected to penetrate through
porous bone and cracks forming irregular shapes depending on the morphology of fractures
]. Hence, the injected cement may be subjected to various stresses including torsion,
flexion, and compression. A finite element analysis demonstrated that the maximum stresses
generated in the injected cement after vertebroplasty may range from 5 to 15 MPa . In the
current study, a representative flexural stress of 10 MPa was used to extrapolate number of
failure cycles thereby allowing comparison of fatigue life amongst materials. The predicted failure
cycle upon applying this flexural stress of experimental composites was comparable to that
commercial products (~ 108 cycles). This may ensure a long-term mechanical performance of
experimental bone composites.
Replacing diluent TEGDMA by PPGDMA provided beneficial effects such as increased
inhibition time, increased final monomer conversion, and decreased calculated polymerization
shrinkage and heat generation for the experimental bone composites. PPGDMA also
promoted hygroscopic expansion to compensate polymerization shrinkage and enhanced
strontium release. Additionally, no detrimental effect on mechanical properties of the composites
was observed upon replacing TEGDMA by PPGDMA. Increasing MCPM level enhanced
hygroscopic expansion, surface apatite formation, and strontium release. Increasing these
reactive fillers reduced static strength of the composites but did not significantly reduce fatigue
resistance of the composites.
S1 File. Raw data. Experimental and commercial bone composites raw data.
DMG has supplied monomers and fillers. SULZER supplied syringes and mixing tips. Dr.
Graham Palmer, Dr. George Georgio, and Dr. Nicola Mordan provided technical support.
Conceptualization: Muhammad Adnan Khan, Kirsty Main, Anne Margaret Young.
Data curation: Piyaphong Panpisut, Kirsty Main, Mayda Arshad, Anne Margaret Young.
Formal analysis: Piyaphong Panpisut, Anne Margaret Young.
21 / 25
Funding acquisition: Piyaphong Panpisut, Kirsty Main, Wendy Xia, Anne Margaret Young.
Investigation: Piyaphong Panpisut, Mayda Arshad, Anne Margaret Young.
Methodology: Piyaphong Panpisut, Muhammad Adnan Khan, Kirsty Main, Anne Margaret
Project administration: Piyaphong Panpisut, Wendy Xia, Anne Margaret Young.
Resources: Wendy Xia, Anne Margaret Young.
Supervision: Wendy Xia, Haralampos Petridis, Anne Margaret Young.
Validation: Anne Margaret Young.
Visualization: Piyaphong Panpisut.
Writing ? original draft: Piyaphong Panpisut, Haralampos Petridis, Anne Margaret Young.
Writing ? review & editing: Piyaphong Panpisut, Haralampos Petridis, Anne Margaret
22 / 25
23 / 25
British Standard. BS ISO 23317:2012 Implants for surgery. In vitro evaluation for apatite-forming ability
of implant materials. Switzerland: BSI Standards Limited; 2012.
Finan L, Palin WM, Moskwa N, McGinley EL, Fleming GJ. The influence of irradiation potential on the
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Mater. 2013; 29(8):906?12. https://doi.org/10.1016/j.dental.2013.05.008 PMID: 23787035
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