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MgCHA particles dispersion in porous PCL scaffolds:
in vitro mineralization and in vivo bone formation
Vincenzo Guarino
1
*
,#
, Silvia Scaglione
2#
, Monica Sandri
3
, Marco A. Alvarez-Perez
1
, Anna Tampieri
3
,
Rodolfo Quarto
4
and Luigi Ambrosio
1
1
National Research Council (CNR) of Italy, Institute of Composite and Biomedical Materials (IMCB), Naples, Italy
2
National Research Council (CNR) of Italy, IEIIT Institute, Genoa, Italy
3
National Research Institute (CNR) of Italy, ISTEC-CNR, Institute of Science and Technology for Ceramic Materials, Faenza, Italy
4
Department of Experimental Medicine (DIMES), University of Genoa, Italy
Abstract
In this work, we focus on the in vitro and in vivo response of composite scaffolds obtained by incorpo-
rating Mg,CO
3
-doped hydroxyapatite (HA) particles in poly(«-caprolactone) (PCL) porous matrices.
After a complete analysis of chemical and physical properties of synthesized particles (i.e. SEM/EDS,
DSC, XRD and FTIR), we demonstrate that the Mg,CO
3
doping influences the surface wettability with
implications upon cell–material interaction and new bone formation mechanisms. In particular, ion
substitution in apatite crystals positively influences the early in vitro cellular response of human mesen-
chymal stem cells (hMSCs), i.e. adhesion andproliferation, and promotes an extensivemineralization of
the scaffold in osteogenic medium, thus conforming to a more faithful reproduction of the native bone
environmentthan undoped HA particles, used as control in PCL matrices. Furthermore, we demonstrate
that Mg,CO
3
-doped HA in PCL scaffolds support the in vivo cellular response by inducing neo-bone
formation as early as 2 months post-implantation, and abundant mature bone tissue at the sixth month,
with a lamellar structure and completely formed bone marrow. Together, these results indicate that Mg
2+
and CO
3
2–
ion substitution in HA particles enhances the scaffold properties, providing the right chemical
signals to combine with morphological requirements (i.e. pore size, shape and interconnectivity) to drive
osteogenic response in scaffold-aided bone regeneration. Copyright © 2012 John Wiley & Sons, Ltd.
Received 10 September 2011; Revised 7 February 2012; Accepted 7 March 2012
Keywords magnesium; hydroxyapatite; ectopic model; mesenchymal stem cells; composite scaffolds;
bone regeneration
1. Introduction
Modern approaches to bone repair involve the use of
bio-inspired scaffolds able to generate new bone tissue with
adequate functional and mechanical properties, and with-
out the risk and expense of using autografts, allografts
and metals (Bruder and Fox, 1999). Different biomaterials
of biological or synthetic origin have been variously
designed to reproduce the functionalities of the bone extra-
cellular matrix (bECM). Recent works have demonstrated
that synthetic polymers with slow degradation rate, such
as poly(e-caprolactone) (PCL), may be efficiently coupled
with apatite particles (Guarino et al., 2008, 2009; Scaglione
et al., 2010) to form composites with good potential for
hard tissue regeneration. The slow degradation rate of
PCL, highly tunable by adjusting composition and molecu-
lar weights, is complementary to the new bone formation
rate, assuring mechanical strength of the order of that of
trabecular bone (Hutmacher, 2001). In particular, the
continuous degradation of the polymer matrix assures
gradual load transfer to the healing tissue, preventing the
stress-shielding atrophy with the stimulation of the healing
and the bone remodelling (De Santis et al., 2009; Ishaug-
Riley et al., 1997). In comparison with other polymers,
e.g. PLGA (Hutmacher, 2001), PCL also shows a good
tendency to support cell growth (Marra et al., 1999),
although recent studies clearly underline its inability to pro-
vide the osteoinductive and osteoconductive cues necessary
* Correspondence to: Vincenzo Guarino, National Research
Council (CNR) of Italy, Institute of Composite and Biomedical
Materials (IMCB), P. le Tecchio 80, 80125 Naples, Italy. E-mail:
vguarino@unina.it
#
These authors contributed equally to this study.
Copyright © 2012 John Wiley & Sons, Ltd.
JOURNAL OF TISSUE ENGINEERING AND REGENERATIVE MEDICINE RESEARCH ARTICLE
J Tissue Eng Regen Med 2014; 8:291–303.
Published online 22 June 2012 in Wiley Online Library (wileyonlinelibrary.com) DOI: 10.1002/term.1521
to guide the cellular processes that underpin the genesis of
new bone tissue (Guarino et al., 2009; Scaglione et al.,
2009). Moreover, PCL shows a detrimental low hydrophili-
city that may hinder the adhesion and colonization of bone
cells and, in turn, the bone regeneration ability of the
scaffold (Amato et al., 2007).
The integration of calcium phosphate particles, such as
hydroxyapatite (HA), within the PCL matrix currently
represents one of the most effective strategies to make
PCL implants more ‘biologically informative’. Indeed,
calcium phosphates have been traditionally used in a
range of orthopaedic and dental applications (Marcacci
et al., 1999) and, more recently, have been investigated
for use as bone tissue-engineering scaffolds because of
their inherent bioactivity (ability to form chemical bond
with bone; Habibovic et al., 2008) and osteoconductivity
(capability of supporting bone growth; LeGeros, 2002).
In the composite scaffolds, ceramic fillers act also as a
reinforcement system able significantly to improve mechan-
ical properties (Guarino et al., 2008). Also, as a binder, PCL
matrix fulfils a protective function for ceramic particles
against non-brittle failure and prevents problems associated
with brittleness and difficulties in shaping hard ceramic
materials to fit bone defects (Coombes and Meikle, 1994).
However, it can often impair bioactive potential by reducing
particle exposure at the interface (Rizzi et al., 2001).
In this context, other calcium phosphates particles (not
only hydroxyapatite) may be considered as potential bio-
active signals for PCL scaffolds. Besides, the chemical
composition of native hydroxyapatite, the main mineral
component of bone tissue and teeth, differs from that of
synthetic hydroxyapatite, due to the presence of several
ionic substitutions in the 3D crystal (i.e. CO
3
2–
,F
–
,Mg
2+
and Na
+
), which play an important role in the biological
responses of bone cells as a function of their spatial distri-
bution and their concentration into the tissue (Elliott,
1994). For example, many authors have demonstrated
that carbonates have a strong influence on the growth of
apatite crystals (Ikoma et al., 2001), sodium plays a role
in bone remodelling (Heaney, 2006) and the fluoride ions
prevent the development of dental caries (Featherstone,
1999). Recently, other authors have demonstrated that
the presence of metal ions (i.e. Mg
2+
,Zn
2+
,Sr
2+
)is
essential to assure a stimulatory effect on bone formation
in vitro and in vivo. In particular, magnesium-doped HA
particles promote osteoblast function, actively participate
in bone regeneration (Dasgupta et al., 2010) and play a
key role in bone metabolism. This effect is mainly evident
during the early stages of osteogenesis, with the stimula-
tion of osteoblast proliferation (Bigi et al., 1992), due to
the enhanced osteoconductivity and resorption of
ion-doped particles in comparison to stoichiometric HA
(Landi et al., 2008; Rude and Gruber, 2004).
In this study we focused on the in vitro and in vivo
cellular response of composite scaffolds incorporating
Mg- and CO
3
-doped apatite particles into PCL matrices.
We demonstrated that Mg,CO
3
doping improves the bio-
logical recognition of porous scaffolds by the extended
study of the in vitro response in osteogenic environment
and in vivo response in an ectopic model. In this context,
we verified that Mg
2+
ions affect some relevant physical
scaffold properties, such as surface wettability or apatite
spatial distribution, which contribute to the promotion
mechanisms involved in bone formation.
2. Materials and methods
2.1. Chemical synthesis of Mg,CO
3
-doped
hydroxyapatite
Mg,CO
3
-doped hydroxyapatite (MgCHA) particles were
synthesized by ISTEC-CNR (Faenza, Italy) and developed
through a neutralization process, based on a simulta-
neously dropwise addition of 600 ml of an acid aqueous
solution containing 88.8 g H
3
PO
4
(Aldrich, 85% pure)
and 200 ml 0.8 MNaHCO
3
(Merck, A.C.S., ISO) aqueous
solution into a basic suspension at 40C for 3 h, prepared
by a dispersion of 100 g Ca(OH)
2
(Aldrich, 95% pure) into
1000mlwater,addedwith48.4gMgCl
2
6H
2
O(Merck,A.C.S.,
ISO) in order to achieve a molar ratio, X
Mg
(Mg:Ca) = 5% in
the mineral phase. The starting Ca:P molar ratio was
therefore equal to the stoichiometric value of HA, 1.67.
The precipitation product was aged for 24 h at 25C,
washed and filtered three times, lyophilized and finally
sieved at 150 mm.
Moreover, HA particles (HAP205, theoretical
density 3.16 g/cm
3
) with bimodal size distribution
(Φ
range
=4–11 mm, Φ
0.5
= 4.02 mm), measured via laser
diffractometry (Sanginario et al., 2006), were supplied by
Plasma Biotal (Tideswell, UK).
2.2. Chemical/physical characterization
2.2.1. Powder composition and physical features
The composition of HA and MgCHA powders was analysed
by SEM/EDS/mapping tests. SEM was preliminarily
performed using a Quanta FEG200 (FEI, The Netherlands)
microscope on dehydrated powders after mild drying (24 h
at 100C). The recognition of elemental constituents,
calcium, phosphorus and magnesium, respectively, in the
powder and their spatial distribution map (mapping test)
were assessed using X-ray energy dispersive spectroscopy
(EDS, Inca200), supported by dedicated software.
Inductively coupled plasma-optical emission spectrome-
try (ICP-OES) analysis (Liberty 200, Varian, Clayton South,
Australia) and Fourier transformed infrared (FTIR)
spectroscopy (Thermo Nicolet-Avatar 320 FTIR) were also
performed to determine HA stoichiometry deviations. The
elemental analyser (LECO C/S, Leco Corp., St. Joseph, MI,
USA) as well as simultaneous thermal analysis (STA 409
Netzsch Geraetebau GmbH, Selb, Germany) were used to
estimate the carbonate amount of the powders. X-ray
diffraction (Cu Karadiation; Miniflex Rigaku, Tokyo,
Japan) was performed to evaluatethe degree of crystallinity,
292 V. Guarino et al.
Copyright © 2012 John Wiley & Sons, Ltd. J Tissue Eng Regen Med 2014; 8:291–303.
DOI: 10.1002/term
the crystalline phase composition and the carbonation
degree of the powders. RX sedimentography (Sedigraph 5100,
Micromeritics, Norcross, GA, USA) was performed in order to
measure the particle size distribution.
2.2.2. Contact angle measurements
Surface wettability was evaluated by contact angle (CA)
measurements performed on PCL films variously incorpo-
rating HA and MgCHA particles. Cast films, obtained by
solvent evaporation, were used to minimize the contribu-
tion of scaffold porosity of the proposed measure. HA or
MgCHA particles, 13% volume fraction with respect to
PCL, were loaded into PCL solution (20% wt in THF),
which was then poured into a specially designed Teflon
mould and kept overnight under the fume hood to slowly
induce solvent evaporation. The samples were classified
according to the surface that had been exposed to air
during casting. Finally, circular discs (10 mm diameter)
were cut with a scalpel blade for wettability tests. The tests
were performed on a Contact Angle (CA) System OCA20
(Dataphysics, Italy) and the CA determined using a tangent
placed at the intersection of the liquid and solid phases by
Software SCA20 (Dataphysics, Italy). For each sample, an
ultrapure water droplet with a volume of 5 ml was dispensed
over six different areas. Measurements were performed at
different times, from 0 to 100 s, to evaluate the influence
of subsequent perfusion flow through the substrate. Contact
angle size was reported as mean standard deviation (SD).
2.3. Bioactive composite design
2.3.1. Preparation of porous substrate
Polycaprolactone-based composite scaffolds were obtained
by the integration of phase inversion and a salt-leaching
technique, as described elsewhere (Guarino et al., 2008).
In brief, poly(e-caprolactone) (PCL; MW 65 kDa) was
dissolved in a tetrahydrofuran (THF; Sigma-Aldrich) to
form a homogeneous solution (20% w/w), gently stirring
for about 3 h at 58C. All chemicals were used as received
from product suppliers. Sodium chloride particles, sieved
into specificsizeranges(212–300 mm) with 90:10 salt:poly-
mer volume ratio were premixed with HA or MgCHA, in
powder form, with the 13% volume fraction with respect
to the polymer volume ratio, and together add to the
polymer solution to reach a homogeneous mixture. The
mixture was placed in an anti-adhesive mould to confer
the cylindrical shape required for the biological tests (diam-
eter 5 mm height 3 mm). Hence, ethanol was used for 24 h
to extract the used solvent and bidistilled water for 7 days to
leach out salt crystals and any other contaminants.
2.3.2. Morphological investigation: SEM supported
by 2D/3D image analyses
The morphology of composite scaffolds was studied by
scanning electron microscopy (SEM; QuantaFEG 200,
FEI, The Netherlands) under high-vacuum conditions
(ca. 10
5
Mbar), using an accelerating voltage of 14 kV.
To improve the sample conductivity, scaffolds received a
preliminary coating of a Pd–Au nanolayer, using a
sputter-coater (Emitech K550, Italy). Moreover, the
samples were scanned using a SkyScan 1072 micro-CT desk
scanner supported by micro-tomographic reconstruction of
cross-sectional images by dedicated software, as reported
elsewhere (Guarino et al., 2010). Image analysis on individ-
ual two-dimensional (2D) slices, previously converted from
cross-sectional images and elaborated by dedicated
software (T-View, Sky Scan, Belgium), permitted the
distinction of single radiopaque phases through differences
in contrast of local regions, offering an estimation of the
spatial distribution of inclusions and reinforcing elements
within the polymer matrix. These images were also elabo-
rated by 2D image analysis via ImageJ v 1.38b (NIH
Freeware; Bethesda, MD, USA) to evaluate the particle
size distribution by the calculation of the particle surface
area (PSA).
2.4. In vitro studies
2.4.1. Cell culture
Biological assays were performed using human mesenchy-
mal stem cells line (hMSCs) obtained from LONZA.
hMSCs were cultured in 75 cm
2
cell culture flasks in
Eagle’sa-minimum essential medium (a-MEM) supple-
mented with 10% fetal bovine serum (FBS), antibiotic
solution (streptomycin 100 mg/ml and penicillin 100 U/ml;
Sigma) and 2 mML-glutamine. For in vitro osteogenic differ-
entiation, hMSC cells were cultured in osteogenic induction
medium consisting of complete a-MEM medium supple-
mented with 50 mg/ml ascorbic acid, 10 mMglycerol-2-
phosphate and 10
–7
Mdexamethasone. hMSCs from
passage 3 were used for all the experimental procedures
and incubated at 37C in a humidified atmosphere (5%
CO
2
, 95% air).
2.4.2. Cell adhesion
The cell adhesion of hMSCs onto PCL, HA and MgHA
scaffolds was evaluated using Vibrant Cell Adhesion Assay
Kit (Molecular Probes). Briefly, hMSCs cultured in a
75 cm
2
cell culture flask were washed with PBS and incu-
bated with calcein AM stock solution to a final concentra-
tion of 5 mMin serum-free medium for 30 min. Af ter
incubation, the cells were washed with PBS, trypsinized
and the cell pellet collected and diluted with culture
medium to obtain the required cell concentration. hMSCs
were seeded onto PCL, HA and MgHA scaffolds and
incubated for 4 and 24 h. Green fluorescence was quantified
using a Wallac Victor
3
1420 spectrophotometer (Perkin-
Elmer, Boston, MA, USA) equipped with 485 nm excitation
and 535 nm emission filters. Percentages of cell adhesion
were obtained by dividing the corrected (background
subtracted) fluorescence of adherent cells by the total
MgCO
3
-doped HA–PCL scaffolds 293
Copyright © 2012 John Wiley & Sons, Ltd. J Tissue Eng Regen Med 2014; 8:291–303.
DOI: 10.1002/term
corrected fluorescence of control cells and multiplying by
100%. Conventional polystyrene 24-well culture plates
were used as a control.
2.4.3. Cell viability
Cell viability of hMSCs plated at a concentration of 1 10
4
in triplicate onto PCL, HA and MgHA composite scaffolds
were checked by the MTT assay for 2, 4 and 6days of
culture. This assay is based on the ability of mitochondrial
dehydrogenases of living cells to oxidize a tetrazolium salt
[3-(4,5-dimethylthiazolyl-2-y)-2,5-diphenyltetrazolium
bromide] to an insoluble blue formazan product. The
concentration of the blue formazan product is directly
proportional to the number of metabolically active cells.
hMSCs seeded onto PCL, PCL–HA and and PL–MgCHA
scaffolds at prescribed times were washed with PBS and
incubated with fresh culture medium containing
0.5 mg/ml MTT for 4 h at 37C in the dark. The superna-
tant was then removed and dimethyl sulphoxide (DMSO)
was added to each well. After 60 min of slow shaking, the
absorbance was quantified by spectrophotometry at
570 nm, using a plate reader. The culture medium during
thw experimental time was changed every other day.
2.4.4. Alizarin red staining
To determine osteogenesis mineralization related to calcium
content of hMSCs cultured on PCL at 7 and 14days, PCL–HA
and PCL–MgCHA composite scaffolds were investigated
using an osteogenesis quantization assay kit (Millipore),
basedonalizarinredS(ARS)staining,whichselectively
binds calcium salts. All scaffold typologies were washed
three times with PBS and fixed with 4% formaldehyde for
1h, then carefully washed five times with distilled water
and stained with ARS (40 mM) for 30 min at room tempera-
ture. After extensive rinsing with distilled water to remove
excess dye, the scaffolds were observed under an optical
microscope. For quantitative analysis of ARS staining, the
PCL, PCL–HA and PCL–MgCHA scaffolds were transferred
to a 1.5 ml microcentrifuge tube and ARS was desorbed with
the use of 10% acetic acid, incubated for 30 min with vortex
andthenheatedat85
C for 10 min. After that, the tubes
werecooledfor5minoniceandcentrifugedat20000g
for 15 min, the pH neutralized with 10% ammonium
hydroxide. Dye sample solution was finally read at 405 nm
in a spectrophotometer. Staining of scaffolds without cells
wasusedasacontrol.TheconcentrationofARSwasdeter-
mined by correlating the absorbance of the experimental
samples with the subtracted value of control scaffold with
a standard curve of known ARS concentrations.
2.4.5. Statistical analysis
All numerical data are presented as mean SD. All
in vitro results were subjected to statistical evaluation
using unpaired Student’st-test to determine significant
differences between groups. The significance level was
set at p<0.05.
2.5. In vivo implant in ectopic model
2.5.1. Cell culture
Bone marrow-derived mesenchymal stem cells (MSCs)
were obtained from the iliac crest of adult sheep under
total anaesthesia. Animal care and surgical procedures were
approved by the Technical Scientific and Ethical Committee
of the Experimental Centre. Bone marrow aspirate was
diluted 1:3 with phosphate-buffered saline (PBS); derived
nucleated cells were isolated using a Ficoll density gradient
solution and counted with a standard nuclear stain (methy-
lene blue). Cells were suspended in Dulbecco’sminimal
essential medium (DMEM) supplemented with 10% FCS,
100 IU/ml penicillin and 100 mg/ml streptomycin and
plated at a density of 1 10
6
cells/cm
2
. The medium was
changed 2days after the original plating and then twice a
week. When culture dishes were nearly confluent, MSCs
were detached with 0.05% trypsin–0.01% EDTA and
replaced until the next confluence. An amount of 1.5–2
million sheep MSCs were suspended in 100 mlculture
medium and statically seeded onto each scaffold (cylinders
of 5 mm diameter 4mm height).
2.5.2. Ectopic bone formation assay in
immunodeficient mice
Bone formation was evaluated by using an established
model of ectopic bone formation (Goshima et al., 1991).
Briefly, an amount of 1.5–210
6
MSCs was combined
with each biomaterial sample and subcutaneously
implanted in immunodeficient mice (CD-1 nu/nu). In
agreement with and the approval of the competent ethical
committee and legal authorities, recipient ID mice aged
1 month, purchased from Charles River Italia (Calco, Italy),
were keptin a controlled environment and given free access
to food and water. The mice were anaesthetized by intra-
muscular injection of xilazine (20 mg/ml) and ketamine
(30 mg/ml). Composites were implanted subcutaneously
on the back of mice (up to four implants for each animal).
The animals were sacrificed 2 and 6 months after
implantation. For each material, at least two constructs
for each experiment were implanted in separate mice;
two independent experiments were performed for a total
amount of four to six samples for each material, for each
time point. The grafts were harvested and processed for
histological analysis.
After sacrificing the animals at the planned experimental
times, the samples were retrieved and fixed in 4% buffered
formalin for 4 h, decalcified with Osteodec (Bio Optica,
Milan, Italy) at 37C for 6 h and dehydrated in an ethanol
series for a total of 6 h. The polymer phase was removed
by dissolution in chloroform for 1 day. The samples were
then paraffin-embedded, sectioned (5 mm thick) at differ-
ent levels, and stained with both haematoxylin and eosin
(H&E) and Masson–Goldner staining. For each scaffold
and bone matrix, blood vessels and mesenchymal tissue
were evaluated.
294 V. Guarino et al.
Copyright © 2012 John Wiley & Sons, Ltd. J Tissue Eng Regen Med 2014; 8:291–303.
DOI: 10.1002/term
The amount of newly formed bone was assessed using
an Olympus BX51 optical fluorescence microscope and
quantified using the free image analysis software ImageJ.
For each scaffold, eight samples (four at each time point)
were harvested after 2 and 6 months of implant. For each
construct, sections 5 mm thick, cut at different levels, were
H&E-stained with for newly formed bone tissue detection.
All images were acquired using both transmitted and
fluorescence light. Images were elaborated by 2D image
analysis, via Image J software. For each construct, the
amount of bone was assessed as percentage of the total
bone tissue vs the total analysed area (bone/area), as
previously described (Martin et al., 2002). Percentages
of new formed bone matrix were presented as mean
standard error (SE). Statistical analysis of the data was
performed using the non-parametric Mann–Whitney test.
3. Results
3.1. MgCHA powder characterization
A preliminary investigation of MgCHA and HA powder
morphology by SEM was assessed to determine the repro-
ducibility of MgCHA synthesis process in terms of final
size and particle shape (Figure 1A, B). SEM images show
relevant differences between the two powder populations.
Figure 1. SEM images and EDS spectra of (A) Mg,CO
3
-doped and (B) undoped HA particles. Mapping of chemical elements (right)
shows respectively the spatial density of P (red dots), Ca (green dots) and Mg (blue dots) elements and their overlapping
MgCO
3
-doped HA–PCL scaffolds 295
Copyright © 2012 John Wiley & Sons, Ltd. J Tissue Eng Regen Med 2014; 8:291–303.
DOI: 10.1002/term
HA particles with micrometric size showed a rounded
shape and well-defined boundaries and without agglom-
eration, in agreement with the indications of the produc-
ing company. In contrast, MgCHA particles showed a wide
distribution of sizes (micro- and submicrometric scale),
which is partially due to a higher tendency to form
clusters (Figure 2A). Sedimentography analyses performed
in the case of MgCHA nanoparticles showed a particle size
distribution of Φ
range
=0.7–1mm (Figure 2B). The results
obtained from the sedimentation method should be consid-
ered as agglomerate size rather than particle size; the
powder is characterized by larger agglomerates formed by
smaller primary particles of diameter ca. 30–50 nm
(Figure 2C). The EDS analyses also confirmed the expected
stoichiometric ratio of the apatitic phases. Two large peaks
due to Ca and P elements were observed in the case of HA
particles with a Ca:P atomic ratio of 1.65, i.e. close to the
theoretical value of stoichiometric HA (Ca:P = 1.67). A
clear peak of magnesium was detected in the EDS spectrum
(ca. 5% with respect to Ca and P elements) from MgCHA
powder. In this case, a (Ca + Mg):P ratio of 1.65 was
detected, which proves a finely controlled Mg substitution
in Ca ion sites of the HA crystal ones during the synthesis
process. Moreover, elemental distribution mapping of
single elements described the spatial distribution of
elements in the powder, clearly showing a strict coexistence
of green, red and blue dots –labels of Ca, P and Mg
elements, respectively, as reported in the merged map.
Moreover, ICP-AES analysis measured Mg content
corresponding to ca. 30–40% of the total amount introduced
in the starting solutions, pointing out the realization of a Mg
doping, ca. 6.8mol.% with respect to Ca, corresponding to
ca. 1.2 wt% of the powder, i.e. close to the Mg biological
content. The (Ca + Mg):P molar ratio values am ounted to
1.85, confirming the carbonation of apatite in the B position
(CO
3
2–
group enters into the crystalline structure of the
apatite substituting PO
4
3–
groups).
The low temperature of the synthesis and the presence
of both carbonate and magnesium ions had a synergistic
effect, making the apatite nuclei smaller and inhibiting
the crystallization in the reaction site of the MgCHA
powder compared to stoichiometric HA. XRD data in fact
showed that the powder was formed by pure apatitic phase
with crystallinity extent Xc (calculated by an experimental
method reported elsewhere; Landi et al., 2000) decreased
as consequence of Mg,CO
3
doping (Figure 3A). FTIR analy-
sis (Figure 3B) also confirmed the crystallinity trend, giving
spectra with a broadened profile. Together with the
phosphate bands at 980–1100 and 560–600 cm
1
, the
presence of both the absorbed and the occluded water
was detected, being respectively referred to the broad band
around 3500 cm
1
and to the peak at 1640 cm
1
.TheFTIR
analysis also showed (Figure 3B) that, thanks to the
selected synthetic procedure, the carbonation spontaneously
occurred in the B position (substitution in the phosphate site:
CO
3
2
stretching signals at 1410 and 1450 cm
1
and bending
peak at 873 cm
1
, referring to B-type phosphate site).
TGA (Figure 4) estimated a carbonation extent of about
6% wt, which the FTIR analysis (Figure 3B) showed to
be preferentially in the B site. The carbonation extent is
responsible for (and in agreement with) the increase of
the (Ca + Mg):P molar ratio from the HA stoichiometric
value 1.67–1.85 (value found by ICP for MgCHA; Figure 2).
The weight loss detected by TGA at high (>1100C) is due
to the well-known dehydroxylation process.
3.2. Effects of Mg,CO
3
doping on scaffold
properties
The influence of the intrinsic chemical signal of Mg ions
and the spatial distribution of MgCHA particles on the
hydrophobic behaviour of PCL porous architecture was
investigated. Contact angle measurements were performed
Figure 2. Particle size distribution of the MgCHA powder determined by Sedigraph
296 V. Guarino et al.
Copyright © 2012 John Wiley & Sons, Ltd. J Tissue Eng Regen Med 2014; 8:291–303.
DOI: 10.1002/term
onto cast films to neglect the effect of porosity on the
droplet formation and shaping. Contact angles at zero time
varied slightly from 87.1 1.8to 83 4.2moving from
HA–PCL to MgCHA–PCL substrates (Figure 5A, B). After
100 s of droplet residence onto the surface, similar trends
were detected, with an angle reduction from 75.42.4
to 69.4 6.3in the case of MgCHA–PCL films. It is note-
worthy that particle clustering, and not homogeneous
spatial distribution, mainly occurs in MgCHA–PCL films,
thus slightly altering the wettability measurement, as
confirmed by values of contact angles with high standard
deviations (SDs) (Figure 5B). This preferential tendency of
MgCHA particles to organize themselves in particles agglom-
erates was confirmed by SEM images of composite scaffolds
(Figure 6). Although no remarkable differences may be
detected in terms of pore architecture (Figure 6A–E), viz.
porosity and pore size (data not shown) as a function of the
used filler type, SEM images at higher magnifications
(Figure 6B–F) clearly show a better dispersion of HA particles
in the PCL porous matrix, while evident clusters of MgCHA
particles may be detected on the pore surfaces. In both cases,
particles partially cover the pore surface, so easily ehxibiting
them outside. Slices from the 3D scaffold taken from
microCT reconstruction qualitatively confirmed the presence
of a more homogeneous population of HA particles of
micrometricsizealongtheporestruts(Figure6C).
Figure 3. XRD (A) and FTIR (B) analyses of Mg,CO
3
-doped HA powder
Figure 4. TG-DSC analyses of Mg,CO
3
-doped HA powder
Figure 5. Effect of Mg and CO
3
doping on wettability of contact angle measurements on MgCHA–PCL and HA–PCL cast films
MgCO
3
-doped HA–PCL scaffolds 297
Copyright © 2012 John Wiley & Sons, Ltd. J Tissue Eng Regen Med 2014; 8:291–303.
DOI: 10.1002/term
Figure 6. SEM images and image analysis of pore architectures of (A–D) HA–PCL, (E–H) MgCHA–PCL scaffolds
298 V. Guarino et al.
Copyright © 2012 John Wiley & Sons, Ltd. J Tissue Eng Regen Med 2014; 8:291–303.
DOI: 10.1002/term
Contrariwise, red domains associated with MgCHA
clusters (Figure 6G), with sizes of several tens of mm,
were recognized. These results were further quantita-
tively validated by PSA distribution calculation via image
analysis (Figure 6D–H).
3.3. In vitro studies
The biocompatibility of the PCL, HA–PCL and MgCHA–PCL
composite scaffolds was studied in terms of cell adhesion
and viability response of hMSCs culture in osteogenic
medium. Figure 7A shows the percentage of attached cells
on the composite surfaces, calculated using the MTT test.
In this case, it has been assumed that the number of cells
in tissue polystyrene plate control, after 4 and 24h of incu-
bation, corresponds to 100% attachment. For MgCHA–PCL
composite scaffolds, the number of adhering cells increases
drastically, attaining 65% after 24 h of cell culture, whereas
the cell number remains around 50% in the case of HA–PCL
scaffolds. Figure 7B also displays differences in cell prolifer-
ation onto undoped HA–PCL and MgCHA–PCL composites.
Higher absorbance values were found in the presence of
MgCHA, thus proposing Mg,CO
3
-doped scaffolds as the
best substrates for cell adhesion and proliferation.
Figure 8 shows the quantitative mineralization assay by
alizarin red staining at time 7 and 14 days of culture in all
cases. To analyse the effective mineralized extracellular
matrix deposition by hMSCs under osteogenic induction, a
semiquantitative analysis of ARS staining was performed.
A calcium content assay revealed higher amounts of Ca
deposits in the case of MgCHA–PCL scaffolds, progressively
decreasing in the case of HA–PCL and PCL scaffolds.
However, at 14 days of culture even the PCL scaffolds show
a relevant increase of mineralization, thus indicating the
attitude of scaffold per se to promote a mineral phase
deposition in osteogenic medium.
3.4. In vivo ectopic implants in mice
After 1, 2 and 6 months of ectopic implantation in ID mice,
constructs were harvested and histologically analysed to
evaluate bone tissue formation and vascularization. No
evidence of inflammatory response was observed in any
scaffold (Figure 9).
No bone formation was ever observed in PCL scaffolds
at any time after the implants (data not shown). Two
months post-implantation, bone matrix deposition was
not observed in HA–PCL scaffolds (Figure 9A), while bone
was evident in the MgCHA–PCL scaffolds, with active
osteoblasts (ob) well visible at the interface between
neo-bone tissue and the surrounding undifferentiated
mesenchymal tissue (Figure 9B). After 6 months of
implantation, only a few pores of the HA–PCL scaffolds
were filled by bone matrix (Figure 9C). In contrast, at
the same time point a significantly important amount of
bone was observed in MgCHA–PCL scaffolds, where
mature bone tissue was massively present within the
scaffold. Almost all pores were completely filled by the
neo-deposited bone with presence of bone marrow (bm);
osteocytes (oc) were entrapped inside their lacunae, and
only a few osteoblasts were detectable. Moreover, the bone
tissue displayed a lamellar structure, developed in a regular
centripetal fashion from the wall of the macropores toward
the central bone marrow cavity (Figure 9D).
The amount of bone matrix deposition within the
implants was quantified. Histological analysis and
computer-assisted image analysis revealed that the bone for-
mation within MgCHA–PCL scaffolds reached 21.12 2.42%
of the total area after 2 months and 26.26 3.88% after
6months. In HA–PCL samples, on the other hand, by
6 months newly formed bone was only 7.840.94% of the
total area. Differences between HA–PCL and MgCHA–PCL
samples were statistically different, after both 2 and
6 months of implantation, in terms of percentage bone
matrix deposition (p<0.05).
The newly formed bone was well vascularized; a signifi-
cant number of blood vessels were observed within the
pores, as well as between contiguous pores running
through the interconnections; a detail of a blood vessel
filled with red blood cells and colonizing a single pore of
the MgCHA–PCL scaffold is shown in Figure 10. Osteoblasts
adhering to the surface of the scaffold primed the bone (b)
matrix deposition towards the inner volume of the pore,
Figure 7. Adhesion and proliferation of hMSCs onto MgCHA–PCL and HA–PCL scaffolds. *Statistically different values (p<0.05)
between paired groups of scaffolds at the same time point
MgCO
3
-doped HA–PCL scaffolds 299
Copyright © 2012 John Wiley & Sons, Ltd. J Tissue Eng Regen Med 2014; 8:291–303.
DOI: 10.1002/term
remaining buried inside their lacunae by the matrix itself,
possibly evolving towards osteocytes (oc). The newly
formed matrix displayed a high level of 3D organization.
A well-organized matrix of collagen fibres was observed,
starting along the pore walls and gradually becoming
concentric to the inner bone marrow cavity. At the same
time, a significant amount of collagen fibres were secreted
by the progenitor cells and organized inside the microporos-
ities of the PCL-based scaffold, just beyond the material–
bone interface.
Histological evidence of osteoconduction of the
MgCHA–PCL material was evident at all time points. In
particular, at 6 months (Figure 11), areas of newly formed
bone (b), bone marrow (bm) and implant material
(i.e. PCL) were observed in direct continuity with PCL
trabeculae completely embedded in the neo-bone tissue.
4. Discussion
The in vivo performance of a bone scaffold is determined
by its ability to stimulate either the onset or completion
of bone defect repair. Accordingly, a wide set of material
properties is required to assure successful post-implant
performance of the device. Basically, a scaffold must be
biodegradable and, preferably, porous. Biodegradability
is mandatory in order to obtain full conversion of the
scaffold into mature and mechanically viable bone, whilst
the presence of tailored pores facilitates the acceleration
of this process (George et al., 2010). However, several
investigations have highlighted that morphologically
designed scaffolds made of synthetic biodegradable poly-
mers show significant deficiencies, due to limitations in
both mechanical properties and in their capacity to be
recognized by cells, both in in vitro and in vivo (Engler
et al., 2006; Mistry et al., 2010). Indeed, scaffold gradually
changes pore architecture and mechanical response as the
degradation mechanisms occur, releasing low molecular
weight products able to variously affect the in vitro and
in vivo response (Bohner et al., 2011). Too rapid degrada-
tion of bone substitutes will often result in the disordering
of osteogenic activities and the dominance of connective
Figure 8. Mineralization by alizarin red staining: quantitative
assay of mineral deposits in PCL, HA–PCL and MgCHA–PCL
scaffolds at 7 and 14 days. *Statistically different values (p<0.05)
between paired groups of scaffolds at the same time point
Figure 9. In vivo bone formation after 2 (A, B) and 6 (C, D) months of implantation within HA–PCL (A–C) and MgCHA–PCL (B–D);
b, bone matrix; ob, osteoblasts lining cells; oc, osteocytes; bm, bone marrow. H&E stain. Bar = 50 mm
300 V. Guarino et al.
Copyright © 2012 John Wiley & Sons, Ltd. J Tissue Eng Regen Med 2014; 8:291–303.
DOI: 10.1002/term
tissue healing, whereas too slow degradation kinetics will
hinder replacement by natural bone (Liu et al., 2010).
Therefore, the combination of bioactive inorganic particles
and slowly-degradable polymer matrices may lead to faster
anchoring to host tissues after implantation, due to the bio-
active signal of the ceramic phase, preserving the scaffold’s
chemical stability until the new bone is totally formed
(Geffre et al., 2009; Lickorish et al., 2007). Here, we
proposed the investigation of composite scaffolds, a PCL
matrix endowed with HA particles after Mg,CO
3
doping.
The use of a proven scaffold technique, such as phase inver-
sion/salt-leaching (Guarino et al., 2009), assure the
required morphology –multiscale porosities composed of
macro- and micropores of preordered sizes –which assures
the creation of a fully interconnected pore network
needed to support correct fluid transport, cell migration
and, finally, proper vascularization of ingrowth tissue.
In particular, microporosity along the pore walls contri-
butes to enhancement of bone formation in the macro-
pores, thus sustaining nutrient exchange, oxygen
transport and multiscale osteointegration (Lan Levengood
et al., 2010). Also, the presence of pores with average
diameter in the range 200–400 mmmaysupportblood
vessel invasion, and would induce osteoblast migration,
adhesion, proliferation and differentiation inside the pores
(Boyan et al., 1996).
Although recent studies have emphasized the crucial
role of porosity features (i.e. pore size and shape) on
cell–material interactions (Betz et al., 2010), it is recognized
that pore architectures with tailored porosity need to be
complemented by bone-inducing agents to stimulate native
bone-formation activity (Liu et al., 2010). Here we aimed to
investigate how chemical modification of HA crystal signals,
due to Mg
2+
and CO
3
2–
ion substitution, combines with
optimized morphological features (i.e. pore size, surface
curvature, interconnectivity) in the fine-tuning of the bio-
logical response of the composite scaffold, triggering the
osteogenic response under in vitro and in vivo conditions.
First, we demonstrated that the presence of Mg
2+
ion
substitution influences the spatial distribution of particles
and wettability of the pore surface. It is well known that
composition (Webb et al., 2000), topography (Anselme
et al., 2000) and wettability are able to regulate a large
number of biological events, including protein adsorption.
Despite some authors recently reporting that the inclusion
of apatite particles into a PCL matrix did not improve the
Figure 10. A single pore cavity of PCL–MgCHA was displayed infiltrated by new bone matrix (green) and osteocytes (dark brown). A
blood vessel with erythrocytes (orange) running through the pore was also displayed (arrows), using the Masson–Goldstein staining
kit; right panel shows enlarged view of region indicated in left panel. Collagen fibres (arrows) were secreted by the progenitor cells
and deposited inside the micro-porosities of the scaffold (s). Bar =100 mm
Figure 11. Morphological evaluation of the MgCHA–PCL scaffold harvested 6 months after implantation. Staining, Masson–Goldstein
(A) and H&E (B). Traces of PCL were visible within the sample; bone (b) and bone marrow (bm) tissues were evident at high
magnification. Bar = 50 mm
MgCO
3
-doped HA–PCL scaffolds 301
Copyright © 2012 John Wiley & Sons, Ltd. J Tissue Eng Regen Med 2014; 8:291–303.
DOI: 10.1002/term
hydrophilicity of the composite scaffold (Salerno et al.,
2011), their results are somewhat contradictory and related
comments vague, due to enormous variability of calcium
phosphate properties as a function of chemical composition
(i.e. ion doping, Ca:P ratio). Other recent near- and
medium-infrared (NIR and MIR) spectroscopic studies
demonstrated that Mg-doped HA is able to retain more
water at its surface than stoichiometric HA, with water
molecules coordinated to cations and adsorbed in the
multilayer form (Bertinetti et al., 2009). Hence, a
hydrated layer, enriched in Mg
2+
ions, may be postulated
for the Mg-doped HA particles exposed on the inner pore
surface. This is consistent with the activation of specific
ionic interactions, offering an effective biophysical signal
capable of triggering the cell interaction mechanisms
involved in the bone-inspired regeneration processes.
This effect is also indicated by several reports of ce-
ramic–polymer composites favouring cell anchorage and
growth, arising from the ability of calcium phosphate to
absorb several important extracellular matrix proteins
(Midy et al., 2001; Pezzatini et al., 2006; Stayton et al.,
2003). In this work, we clearly show that hMSC attach-
ment and proliferation are influenced by the presence of
calcium phosphates in the polymer matrix, the greatest
improvement being apparent in the case of MgCHA–PCL
scaffolds. Meanwhile, alizarin red quantitative assays
clearlyindicatetheroleofcalciumphosphates,especially
in the case of MgCHA particles, in promoting in vitro
mineralization. These results may be dependent upon
the variation in crystallinity detectable in the different
calcium phosphate forms (Celotti et al., 2006), as
recently described in the literature. In particular, the
presence of magnesium stabilizes amorphous HA, so
preventing its transition to the crystalline form (Mathai
and Shozo, 2001). This is consistent with previous studies
demonstratig that the presence of magnesium promotes
the nucleation of a high number of apatite nuclei (Landi
et al., 2008) but tends to inhibit the crystallization
phenomena, due to the reduction in the Ca:P molar ratio
in comparison with undoped HA (Bigi et al., 1993).
Indeed, the substitution of bivalent ions such as Mg
2+
to the central calcium atom in the stoichiometric HA
structure is accompanied by a slight change of complex
bond energy, improving the chemical stability of HA
(Gutowska et al., 2005). Consequently, the higher
chemical stability and lower crystallinity of MgCHA
should promote the in vitro production of mineral
deposits, in agreement with the studies of Morgan et al.
(1996) showing that lower HA crystallinity results in a
higher amount of mineralization of excreted extracellular
matrix.
In this context, it is worth noting that the greater
tendency of MgCHA to form clusters, compared to HA,
did not reduce the osteogenic effect of particles during
in vitro culture, thus showing that mineralization is
mainly promoted by the presence of specific chemical
signals (i.e. Mg
2+
and CO
3
2–
substitution), and only
secondarily by morphological features (i.e. size and spatial
distribution of particles).
On the other hand, in vivo results highlighted how the
pattern of newly formed bone formed within scaffolds is
deeply influenced by the morphological cues of the grafts
and also by their chemistry, which deeply affects the cells
fate at early stages. In particular, we have demonstrated
here that MgCHA–PCL scaffolds provide an osteomimetic
surface and an efficient template for newly formed bone
tissue. Osteoprogenitor cells initially deposited bone
matrix along the walls of the macropores, recognizing
the available surface as pre-existing natural bone. At the
same time, cells consolidated osteointegration with the
polymeric structure, through a mesh of fibres infiltrating
the micropores of the scaffold. The newly formed bone
tissue was subsequently remodelled over several months,
with mature lamellar bone surrounding a central bone
marrow cavity.
5. Conclusion
A key current strategy in regenerative medicine is the
critical study of natural tissues and biomimesis to inform
the design of new composite materials, able to emulate
the structural and functional response of tissue. In the
case of bone, the mineralized extracellular matrix gener-
ally consists of organic polymers (e.g. collagen, protogly-
cans) in which are dispersed non-stoichiometric apatites
of poor crystallinity, containing cationic and anionic
substitutions at sites within the HA crystal structure.
These polymer-based scaffolds, incorporating synthetic
apatite fillers, are considered as the ‘gold standard’in
scaffold design for bone. In this study, we have demon-
strated that the use of porous scaffolds made of PCL
matrix, endowed with Mg,CO
3
-doped HA particles, facili-
tates the creation of a microenvironment which is amenable
to bone cells, able to trigger osteogenic mechanisms
through a set of chemical and morphological information
intrinsically programmed into the composite materials. In
particular, ionic substitution stimulates cells to produce
mineral extracellular matter through the activation of early
osteogenesis mechanisms. Meanwhile, porous architec-
ture with tailored porosity features (i.e. pore size,
concavity) supports the in vivo growth of mature bone
with hierarchical organization by imparting the natural
osteon-like structure of bone, with lamellae centripetally
assembled from the wall of the macropores toward
the central bone marrow cavity, during the first 6 months
of implantation.
Acknowledgements
This study was supported from the Ministero dell’Universita e
della Ricerca by funds of Rete Nazionale di Ricerca TISSUE-
NET (Grant No. RBPR05RSM2). Scanning electron micros-
copy was supported by the Transmission and Scanning
Electron Microscopy Labs (LAMEST) of the National
Research Council.
302 V. Guarino et al.
Copyright © 2012 John Wiley & Sons, Ltd. J Tissue Eng Regen Med 2014; 8:291–303.
DOI: 10.1002/term
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