Preparation of the fast setting and degrading Ca–Si–Mg cement with
both odontogenesis and angiogenesis differentiation of human
periodontal ligament cells
Yi-Wen Chen a,b, Tuan-Ti Hsu c, Kan Wang d,e, Ming-You Shie b,*
a
Graduate Institute of Clinical Medical Science, China Medical University, Taichung
City, Taiwan
b
3D Printing Medical Research Center, China Medical University Hospital, Taichung
City, Taiwan
c
Institute of Oral Science, Chung Shan Medical University, Taichung City, Taiwan
H. Milton Stewart School of Industrial and Systems Engineering, Georgia Institute
of Technology, Atlanta, GA 30332, USA
d
e
Georgia Tech Manufacturing Institute, Georgia Institute of Technology, Atlanta, GA
30332, USA
Correspondence:
Ming-You Shie, 3D Printing Medical Research Center, China Medical University
Hospital, Taichung City, Taiwan
E-mail: eviltacasi@gmail.com; +886-4-22052121; fax: +886-4-24759065
Acknowledgements
The authors acknowledge receipt of a grant from the Ministry of Science and
Technology grants (MOST 104-2314-B-039-004). The authors declare that they have
no conflicts of interest.
1
ABSTRACT
Develop a fast setting and controllable degrading magnesium-calcium silicate cement
(Mg-CS) by sol-gel, and establish a mechanism using Mg ions to stimulate human
periodontal ligament cells (hPDLs) are two purposes of this study. We have used the
diametral tensile strength measurement to obtain the mechanical strength and stability
of Mg-CS cement; in addition, the cement degradation properties is realized by
measuring the releasing amount of Si and Mg ions in the simulated body fluid. The
other cell characteristics of hPDLs, such as proliferation, differentiation and
mineralization were examined while hPDLs were cultured on specimen surfaces. This
study found out the degradation rate of Mg-CS cements depends on the Mg content in
CS. Regarding in vitro bioactivity; the CS cements were covered with abundant
clusters of apatite spherulites after immersion of 24 h, while less apatite spherulites
were formatted on the Mg-rich cement surfaces. In addition, the authors also explored
the effects of Mg ions on the odontogenesis and angiogenesis differentiation of
hPDLs in comparison with CS cement. The proliferation, alkaline phosphatase,
odontogenesis-related genes (DSPP and DMP-1), and angiogenesis-related protein
(vWF and ang-1) secretion of hPDLs were significantly stimulated when the Mg
content of the specimen was increased. The results in this study suggest that Mg-CS
materials with this modified composition could stimulate hPDLs behavior and can be
good bioceramics for bone substitutes and hard tissue regeneration applications as
they stimulate odontogenesis/angiogenesis.
Keywords: Magnesium; Calcium silicate; Biodegradable; Periodontal ligament cells;
Odontogenic; Angiogenic
2
1. Introduction
Bone substitutes have been shown to have great promising potential for hard
tissue regeneration by stimulating tissue regeneration and forming calcified tissue [1].
Among the variety of bone substitutes, calcium phosphate and calcium silicate have
been most rigorously examined in regards to its benefits for hard tissue formation [26]. Although common β-tricalcium phosphate has been widely used as a bone defect
substitute material due to its favorable desirable biocompatibility, its biodegradation
and osteostimulation properties are still far from optimal [7-9]. In recent years,
several calcium silicate-based (CS) cements have been proved as potential bioactive
materials for use as bone substitute materials; some of them enhance application,
furthermore, some of them perform excellent in vitro and in vivo osteogenesis and
odontogenesis [8,9]. In our previous study, we produced a fast setting CS cement
containing a combination set of compounds of CaO, SiO2, and Al2O3, which was have
shown to decrease setting time of cement [10]. In addition, CS cement not only
exhibits good osteoconduction effects [2,3,5-7,11], but also reduces the inflammation
markers in primary human dental pulp cells (hDPCs) [2,7,12]. Besides, CS based
cements have also owns the ability to stimulate osteogenesis differentiation of various
stem cells, such as bone marrow stromal cells, adipose-derived stem cells, human
dental pulp cells and periodontal ligament cells [8,9,13]. The release concentrations of
Si ions from CS materials affects the behavior of different cell types, such as
inhibiting osteoclastgenesis in macrophage [5], and promoting angiogenesis in hDPCs
[14,15]. However, the amount of Si ions in CS-based materials can affect the
adsorption of various extracellular matrices (ECM) such as collagen I, fibronectin,
and vitronectin, and also promotes the up-regulation up-regulate of MAPK/ERK and
MAPK/p38, signaling the pathway more effectively with higher effectiveness than Ca
3
components [7].
Other studies have found that the incorporation of metal ions into CS ceramics
significantly promote their bioactive characteristic, which has inspired us to consider
the possibility that an alternative approach of bioceramics with modified ions may be
useful as an alternative approach to improving might improve cell behavior for tissue
regeneration. In a previous study, Zhai et al. synthesized and analyzed the properties
of Ca–Si–Mg-containing bioceramics, bredigite (Bre, Ca7MgSi4O16), akermanite (Ake,
Ca2MgSi2O7) and diopside (Dio, CaMgSi2O6). These substances not only showed
excellent characteristics in vitro apatite precipitates, but also led to better mechanical
properties than bioglass [16,17]. In addition, Ca–Si–Mg bioceramics enhance cell
adhesion and osteogenic differentiation osteoblasts, improve human mesenchymal
stem cells and human adipose stem cells by up-regulation of bone-related genes
[16,18]. Due to the advantages of Ca–Si–Mg bioceramics for bone regeneration, it is
was has been recommended that they be are used for to use in periodontal
regeneration by stimulation of cementogenic/angiogenic differentiation of human
periodontal ligament cells (hPDLs) [19]. Recently, several studies have investigated
whether hPDLs constitutively express osteogenesis cytokines and growth factors,
such as COL I, osteopontin, ALP, bone morphogenetic proteins, and antiinflammatory cytokines [20]. Similarly, hPDLs are a potential candidate for
periodontal tissue engineering and have been shown validated to differentiate
osteoblastic, fibroblastic and cementoblastic lineages [21]. However, the mechanism
through which Mg-CS stimulates odontogenic/angiogenic differentiation of hPDLs is
not completely known.
In this study, we will investigate in vitro odontogenic/angiogenesis stimulation
of Mg-CS cement for hPDLs as well as the possible mechanism underlying this
4
process. Therefore, the two interaction of both the Mg-CS cements and their ionic
products with hPDLs is are scientifically systematically analyzed in this study,
including cell proliferation, odontogenesis and angiogenesis differentiation of hPDLs.
In addition Last, this article will discuss how the ions from three different Mg
contents in CS cement affect the behavior of hPDLs in cementogenic and
angiogenesis. study also investigates how the ions from three different Mg contents in
CS cement affect the behavior of hPDLs in cementogenic and angiogenesis.
5
2. Materials and methods
2.1. Preparation of specimens
The method used here for the preparation of CS powder has been has been given
in an earlier paper [7]. In brief, reagent-grade tetraethyl orthosilicate (Sigma-Aldrich,
St. Louis, MO, USA), calcium nitrate tetrahydrate (Showa, Tokyo, Japan), and
magnesium nitrate hexahydrate (Showa) were used as precursors for SiO2, CaO, and
MgO respectively. The catalyst was nitric acid (2M), and absolute ethanol was used
as the solvent. The nominal molar ratios of CaO–SiO2–MgO ranged were listed in
Table 1, and the schematic illustration of these cements preparation was showed in
supply material 1. General sol-gel procedures, the sol solution were aged at 60°C for
1 day and the solvent was vaporize in an oven at 120°C for 1 day. After sintering at
800°C for 2 h with a high-temperature furnace, the sintered powders were then ballmilled for 6 h in ethyl alcohol using a centrifugal ball mill machine (Retsch S 100,
Hann, Germany) and then dried in an oven at 60°C. The cement was prepared by hand
mixing the powder with distilled water with different liquid-to-powder (L/P) ratio,
and the L/P ratio was varied over the range 0.38–0.42 mL/g, depending on the kind of
cement. After being mixed with water, the cements were molded in a Teflod mold
(diameter: 6 mm, height: 3 mm). There was enough cement to fully cover each well of
the 24-well plate (GeneDireX, Las Vegas, NV) to a thickness of 2 mm for cell
experiments. All samples were stored in an incubator at 100% relative humidity and
37°C for 1 day of hydration.
2.2. Setting time and strength
After the powder was mixed with H2O, the cements were placed into a
cylindrical mould and stored in an incubator at 37°C and 100% relative humidity for
6
hydration. The setting time of the cements was tested according to standards set by
the International Standards Organization (ISO) 9917-1. For evaluation of the setting
time, each material was analyzed using Gilmore needles (456.5 g), and the time was
recorded when the needle failed to create a 1-mm deep indentation in three separate
areas. After being taken out of the mould, the specimens were again incubated at
37°C in 100% humidity for 1 day. The diametral tensile strength (DTS) testing was
conducted on an EZ-Test machine (Shimadzu, Kyoto, Japan) at a loading rate of 1
mm/min. The DTS value of the samples was calculated from the relationship DTS =
2P/ πbw, where P is the maximal peak load (Newton), b is the diameter (mm) and w is
the thickness (mm) of the samples. The ten specimens were examined for each of the
materials.
2.3. Phase compositions and morphology
The phase composition of the cements was analyzed using X-ray diffractometry
(XRD; Bruker D8 SSS, Karlsruhe, Germany), run at 30 kV and 30 mA at a scanning
speed of 1o/min. The morphology of the cement specimens were coated with gold and
examined under a scanning electron microscope (SEM; JSM-6700F, JEOL) operated
in the lower secondary electron image (LEI) mode at 3 kV accelerating voltage.
2.4. Weight loss
The degree of degradation was determined by monitoring the weight change of
the specimen following immersion in a simulated body fluid (SBF) solution. After
drying at 60°C, the specimens were weighed both before and after immersion using a
balance (TE214S, Sartorius, Goetingen, Germany). Ten specimens were examined for
each of the materials investigated at each time point.
7
2.5. Ion concentration
After immersion for different time-points, the Ca, P, Si and Mg ions
concentration released from materials on SBF were analyzed using an inductively
coupled plasma-atomic emission spectrometer (ICP-AES; Perkin-Elmer OPT 1MA
3000DV, Shelton, CT, USA). Three samples were measured for each data point. The
results were obtained in triplicate from three separate samples for each test.
2.6. Cell isolation and culture
The hPDLs were freshly derived from caries-free, intact premolars that had been
extracted from 3 healthy adults (18-24 years of age) for orthodontic treatment
purposes. The patient gave informed consent, and approval from the Ethics
Committee of the Chung Shan Medicine University Hospital was obtained (CSMUH
No. CS14117). The teeth were instantly immersed into a phosphate-buffered saline
(PBS; Caisson Laboratories, North Logan, UT, USA) containing 100 U/mL
penicillin/streptomycin (Caisson) and kept immersed while they were transferred to
the laboratory. The PDL tissues of the root surface were separated by blade, washed
several times with PBS, and then cut into cubes. The tissue was broken down using 1
mL of 2 mg/mL type I collagenase (Sigma-Aldrich) and 4 mg/mL dispase (SigmaAldrich) for 30 min in the incubator. After the tissue fragments had been broken down,
they were distributed into plates and cultured in DMEM, supplemented with 20%
fetal bovine serum (FBS; Caisson), 1% penicillin (10,000 U/mL)/streptomycin
(10,000 mg/mL) (PS, Caisson) and kept in a humidified atmosphere with 5% CO2 at
37°C; the medium was changed every 3 days. The primary cells were then subcultured to obtain passage 0 single cell-derived clones (P0), and the passages P4-P6
8
were used for the following in vitro study. The cementogenic differentiation medium
was DMEM supplemented with 10−8 M dexamethasone (Sigma-Aldrich), 0.05 g/L Lascorbic acid (Sigma-Aldrich) and 2.16 g/L glycerol 2-phosphate disodium salt
hydrate (Sigma-Aldrich). The angiogenic differentiation medium was DMEM
supplemented with 2% fetal bovine serum, 1% penicillin/streptomycin, and 50 ng/mL
vascular endothelial growth factor (Prospec, East Brunswick, NJ).
2.7. Cell proliferation
The hPDLs were suspended in a density of 104 cells/ml and were directly seeded
over each specimen at 37 °C in a 5% CO2 atmosphere. After different culturing times,
cell viability was evaluated using the PrestoBlue® assay (Invitrogen, Grand Island,
NY). The reagent PrestoBlue® was used for real-time and repeated monitoring of cell
cytotoxicity, which is determined based on the level of mitochondrial activity. To
explain this process briefly, at the end of the culture period, the medium is discarded
and the wells are washed twice with PBS. Each well is filled with a solution
containing PrestoBlue® and fresh DMEM (1: 9) at 37°C for 30 min. The solution in
each well is then transferred to a new 96-well plate. Plates are read using a multi-well
spectrophotometer (Hitachi, Tokyo, Japan) at 570 nm with a reference wavelength of
600 nm. Cells cultured on a tissue culture plate are used as a control (Ctl). The results
were obtained in triplicate from six separate experiments for each test.
2.8. Cell morphology
After the cells had been seeded for 1, 3, and 7 days, the materials were washed
three times with cold PBS and fixed in 1.5% glutaraldehyde (Sigma) for 2 h, after
which they were dehydrated using a graded ethanol series for 20 min at each
9
concentration and dried with liquid CO2 using a critical point dryer device (LADD
28000; LADD, Williston, VT). The dried specimens were then mounted on stubs,
coated with gold, and viewed using Scanning electron microscopy (JEOL JSM-7401F,
Tokyo, Japan).
2.9. Odontogenesis differentiation
For the detection of odontogenesis-related gene [dentin matrix protein 1 (DMP1) and dentin sialophosphoprotein (DSPP)] and alkaline phosphatase (ALP)
expression of PDL cells, which were cultured at a density of 104 cells per well in
separate 6-well plates. After 24 h of incubation, the culture medium was removed and
replaced by conditioned medium or control medium. Total RNA of all four groups
was extracted using TRIzol reagent (Invitrogen) after 3 days and analyzed by RTqPCR. Total RNA (500 ng) was used for the synthesis of complementary DNA using
cDNA Synthesis Kit (GenedireX) following the manufacturer’s instructions. RTqPCR primers (Table 2) were designed based on cDNA sequences from the NCBI
Sequence database. SYBR Green qPCR Master Mix (Invitrogen) was used for
detection and the target mRNA expressions were assayed on the ABI Step One Plus
real-time PCR system (Applied Biosystems, Foster City, California, USA). Each
sample was performed in triplicate.
2.10. Alizarin Red S Stain
The accumulated calcium deposition was analyzed using Alizarin Red S
staining, following a method developed for a previous study.[22] In brief, the
specimens are fixed with 4% paraformadedyde (Sigma-Aldrich) for 15 min and then
incubated in 0.5% Alizarin Red S (Sigma-Aldrich) at pH of 4.0 for 15 min at room
10
temperature. After this, the cells are washed with PBS and photographs are taken
using an optical microscope (BH2-UMA, Olympus, Tokyo, Japan) equipped with a
digital camera (Nikon, Tokyo, Japan) at 200x magnification. The Alizarin Red is also
quantified using a solution of 20% methanol and 10% acetic acid in water. After 15
min, the liquid is transferred to a 96-well, and the quantity of Alizarin Red is
determined using a spectrophotometer at 450 nm.
2.11. Angiogenic assay
The Matrigel in vitro angiogenesis assay is used (Corning, MA, USA) to
evaluate the angiogenic properties of hPDLs cultured with different materials. After
being cultured for 1 day, the cells are photographed using an inverted microscope
(Olympus, Tokyo, Japan). In addition, the production of ang-1 and vWF were
quantified using ELISA kits (Abcam, catalog no. ab99970 and ab108918) according
to the manufacturer’s instructions. To summarize this process briefly, the hPDLs are
cultured on substrates for 3 and 7 days, and proteins from the whole cell lysates are
then collected and quantified using the ELISA kit.
2.14. Statistical analysis
A one-way variance statistical analysis was used to evaluate the significance of
the differences between the groups in each experiment. Scheffe’s multiple comparison
test was used to determine the significance of the deviations in the data for each
specimen. In all cases, the results were considered statistically significant with a p
value < 0.05.
11
3. Results and discussion
3.1. Setting time and strength of cements
The powder and liquid are mixed in an appropriate ratio to form a paste that
hardens through the interaction of the crystals precipitated in the paste when it is at
body temperature. Thus, the setting time is one of most important factors in the
clinical use of this compound [23]. The long setting time will lead to clinical
problems under certain conditions. We propose that 10–15 min is a suitable setting
time interval in the clinical [24]. The setting time of all cement specimens used in this
study are listed in Table 1. The setting time of the cements increase when the Mg
content is increased; the setting time of the Mg incorporated with different amounts of
calcium silicate, can be ranged from 16 min (CS) to 21 min (Mg10). In the present
study, the setting time of calcium-silicate based cements is proportional to the amount
of calcium silicate hydrate (CSH) [2,25]. CSH is able to decrease the setting time of
the calcium-based cement. Our results show that a setting time of approximately 20
min for injectable bone cements for use in vertebroplasty and kyphoplasty is possible
[4]. The DTS values of the hardened cements are 2.5, 2.4, and 2.2 MPa, with the
values changing in response to increases in the Mg content (Table 1). ANOVA shows
significant differences between the DTS values of the three materials. Thus,
differences in the phase composition of the hydration products may play a crucial role
in the mechanical properties of these cements.
3.2. Phase composition of cements
Figure 1 shows the XRD patterns of the specimens after hydration for 1 day. CS
has an obvious diffraction peak near 2θ = 29.4°, which corresponds to the CSH gel,
and incompletely reacted inorganic component phases of the β-dicalcium silicate (β-
12
Ca2SiO4) at 2θ between 32° and 34° [25]. Contrary to our findings, the appearance of
a poorly crystalline CSH gel was the principal hydration product of Mg-contained
specimens such as Mg5 and Mg10. The specimens containing Mg showed that the
cements are mainly composed of bredigite phase (JCPDS 33- 0399). The broad peak
of the magnesium silicate hydrate (MSH) gel phase seems to overlap with the
incipient presence of a crystalline bredigite phase between 32° and 37° [26].
3.3. Characterization of immersed cements
The DTS values of the as-set cement specimens were 2.5, 2.4, and 2.2 MPa for
CS, Mg5, and Mg10, respectively (Fig. 2). The values of CS after 1-, 4-, and 12weeks of immersion were 2.7, 3.6, and 4.0 MPa, respectively. Interesting, the results
revealed that Mg-contained specimens gradually increased the strength with
increasing immersion time, and thereafter decreased. The time to achieve maximum
strength value decreased as the amount of gelatin increased in the cement powder
mixture, which being in the order: 0 mol% (12 weeks), 5 mol% (4 weeks), and 10
mol% (4 weeks). When soaking for 12 weeks, the strength of the 10 mol% Mgcontaining cement significantly decreased from the maximum value of 3.0 MPa down
to 2.6 MPa (p < 0.05) comparable to the control cement of 4.0 MPa and 5 mol% Mgcontaining cement of 3.0 MPa at the same time. CS is a typical silicate-based cement
that has shown great potential for hard tissue repair applications. It has previously
been demonstrated that diopside ceramics have excellent mechanical strength and the
ability to induce in vitro apatite-formation in SBF [17].
Dissolubility plays an important role in biodegradation, and so must be
considered when trying to develop a material that has a degradation rate most
appropriate to hastening and easing the process of hard tissue regeneration [3]. With
13
this in mind, the degradation rates of the Mg-contained CS specimens in SBF solution
have been recorded for various time-points, as shown in Fig. 3. When soaking in SBF,
all specimens show a decrease in weight with the increase in immersion time, a
decrease of 4–12% after 1 week. Mg5 and Mg10 both have more weight loss than CS.
All the specimens exhibit an increased weight loss with an increase in the immersion
time, and weight losses of 9.5%, 16.9%, and 24.6% were observed for CS, Mg5, and
Mg10, respectively, which indicates a significant difference (p < 0.05) after
immersion for 12 weeks. In clinical use, it is known that bioactive materials with
different degradation rates are required [27]. Kao et al. finds that the degradation rate
of the calcium phosphate/calcium silicate composite cements increase with time when
the calcium phosphate content is more than 20 wt%, and CS cement exhibits only a
small amount of weight loss [25]. Our results exhibit that Mg amounts of CS-based
materials play an important role in determining the degradability of specimens, and
the degradation of cements may be controlled by the adjustment of Mg content.
Moreover, the crystal structure of materials is also affected by the degradation and
mechanical properties of the cements [28]. The fast degradation rate of Mg limits its
applications, especially as a bone substitute [29]. Wu and Chang assert that bredigite
has a higher weight loss than akermanite and diopside after being soaked in SBF for
28 days [30].
Mg is one of the most important ions associated with biological apatite and it
can control biological functions and promote catalytic reaction [31]. The surface
microstructure of the specimens before and after immersion in SBF for 3 and 24 h is
shown in Fig 4. It is readily visible that all specimens exhibit a dense and smooth
surface containing interlocking particles and some micro-pores were evident. After
immersion in SBF for 3 h, the CS cement specimens induced the precipitation of the
14
formation of an apatite spheres. The formation of the bone-like apatite in SBF has
proven to be useful in predicting the bone-bonding ability of material in vitro [32].
For Mg-content specimens, it is evident that the ability of spherical granules
precipitated on the surface of specimens after immersion were decreased when Mgcontent increased. Because Mg is incorporated, the Ca in the center is substitute by it
[33]. Since its ionic radius (0.69 Å) is shorter than Ca (0.99 Å) [34], the substitution
of Mg into Ca position deforms the apatite structure. Therefore, the result shows that
the addition of Mg could not contribute to the tendency to form apatite in SBF.
Besides, the ability of apatite-formed on the specimen appears to be dependent on the
Ca–Si ratio of the calcium silicate based material. The Si–OH functional groups on
the surface of CS-based materials have been shown to act as nucleation centers for
apatite precipitation [25,35].
Variations in the SBF Ca, P, Si, and Mg ion concentrations after being
immersion for various time periods are shown in Fig. 5. For CS, the Ca ion
concentration of the solution was decreased to approximately 1.55 mM by 1 week,
which is lower than the baseline Ca concentration of SBF (p < 0.05). For Mgcontained specimens, the Ca ion concentration of the solution was significant higher
(p < 0.05) than pure calcium silicate cement. As for the P ion concentration of the
solution, it decreased for all time-points. The in vitro bioactivity of the calciumsilicate based materials indicates that the presence of PO43− ions in the composition is
not an essential requirement for the formation of an apatite layer. This is noteworthy
because it is known that PO43− depletes calcium and phosphate ions because the PO43−
ions originate from the in vitro assay solution [36]. The Ca ion was possibly
originating from the less-ordered hydration products, could significantly promote
apatite growth by promoting local Ca supersaturation, as reported in a previous study
15
[8], thereby raising the ionic product of the apatite in the surrounding environment
and promoting the nucleation rate of the apatite. Si concentrations increase steadily
with time. Si ion concentration of solution was 2.41, 2.66, and 2.99 mM were CS,
Mg5, and Mg10 immersed for 12 weeks, respectively. For Mg ion, it is evident that
the concentration released from specimens was increased when Mg-content increased.
CS-based materials undergo hydrolysis on immersion, and Si and Mg ions dissolves
during immersion [36,37]. It is already well-established that Mg ions play an
important role in stimulating cell proliferation, which can promote DNA and protein
synthesis and regulate Mg conducting channels [38].Taking cell functions into
account, the certain amounts of Si released from CS-based materials may affect
desired cell functions [39-41]. The current CS gives the ability to control the rate at
which soluble Si ions promote cell adhesion, proliferation, and differentiation [36,42].
3.4. Cell proliferation
The PrestoBlue® assay was used to quantitatively analyze cell proliferation.
Results confirm that hPDLs can indeed adhered and proliferate on different CS-based
materials (Fig. 6). It was found that hPDLs cultured on CS and Mg-CS have a rapid
growth rate than Ctl (p < 0.05), which increases with culture time. Interestingly,
Mg20 elicits a significant increase, 18% and 20% on days 1 and 3 of being cultured,
respectively, in comparison with CS. Previous studies have shown that increased
concentrations of Mg ions in cell culture media can adversely affect cell proliferation,
and Mg ions act as a mitogenic factor to promote cellular proliferation in vitro.[43,44]
Therefore, if there is indeed an increase in Mg ion concentration in the medium, it is
most likely due to the dissolution of increased Mg-rich amorphous content. In
addition, the numbers of hPDLs adhering on the cements proliferated with the
16
increase of culture time, indicating good biocompatibility of Mg-containing CS
materials [45]. Fig. 7 shows the morphology of hPDLs cultured to various specimens
after different time-points of culture. At day 1, the cells were flat with intact, welldefined morphology and extending filopodia, indicating survival of the hPDLs on all
specimens. After cultured for 3 and 7 days, hPDLs grew well on three kinds of
calcium silicate-based materials, and there appeared to be more cells compared to
those at day 1.
3.5. Odontogenesis
CS-based materials have been shown to induce osteogenic and cementogenic
differentiation in various cell types [14,46]. To investigate the potential of the hPDLs
for odontoblast differentiation after being cultured on Mg-CS materials, the hPDLs
were cultured for up to three days to evaluate the gene and protein expression levels
of differentiation markers using real-time PCR and ELISA. At day 3, Mg10 markedly
up-regulated DMP-1 (Fig. 8A) and DSPP (Fig. 8B), and significance (p < 0.05)
increased 29% and 18% compared with CS, respectively. In ALP, the Mg5 and Mg10
were markedly up regulated that their differences significantly increased (p < 0.05) by
36% and 84%, respectively compared with CS on day 3 (Fig 9A). In addition, to
investigate the effect of different substrates on mineralization in the hPDLs, we
examined mineralized nodule formation in the hPDLs by using alizarin red S staining
(Fig. 9B). Alizarin Red S staining showed data of bone nodule formation and calcium
deposition. The results show that after 7 and 14 days Mg10 clearly increases the area
of calcified nodules compared with CS. There were no significant differences (p >
0.05) between CS and Mg5 after 7 days. On Mg10, there is a significant (1.42-, and
1.44-fold on day 7, and 14, respectively) enhancement (p < 0.05) of calcium content
17
has been observed compared with CS specimens. In previous studies we had shown
that the viability of hDPCs and osteoblast-like cells on CS are higher than Ctl for all
culture times [47]. Shie et al. found that soluble factors from CS-based substrates may
be more important for adhesion and osteogenic differentiation in a growth medium
[22]. The complexity of the bone supporting apparatus makes periodontal tissue
regeneration a challenging field, as it involves the regeneration of cementum, a
functionally oriented PDL tissue and an alveolar bone in the periodontal defect. It has
been demonstrated that hPDLs can be differentiated into various types of cells [48]. In
recent years, several studies have focused on the interaction of biomaterials and
mesenchyme-derived cells and their ability to promote periodontal tissue regeneration
[21]. ALP activity is also associated with hard tissue formation, and is produced in
high levels during the hard tissue formation phase [49]. It was found that Mg5 and
Mg10 materials promote cell proliferation and the differentiation of hPDLs.
Interestingly, Mg-CS cement significantly increases relative odontogenesis gene
(DSPP and DMP-1) and calcium deposition of hPDLs compared to those on the CS
cement. Otherwise, the cells cultured with CS showed OD value higher than Ctl and
this result pointed out the silicon is also an important component element of Mg-CS
cement to facilitate cell proliferation. The previous reports showed that bone
incorporates various nutrients in the form of trace elements and both Si and Mg have
been found to play absolutely vital roles in the bone formation, growth, and
regeneration and they are essential cofactors for enzymes which involve in the
synthesis of the constituents of bone matrix.[50]
3.6. Angiogenesis
18
Cells seeded on Matrigel covered on substrates formed an extensive lattice of
vessel-like structures after 24 hours (Fig. 10A). The hPDLs cultured with Mg5 and
Mg10 increased the size of the junctional area, and the thickness of the vessel-like
structures increased. The vessel-like structures are significantly different than with CS.
ELISA analysis shows that the Mg5 and Mg10 materials significantly promote vWF
(Fig 10B) and ang-1 (Fig 10C) secretion in hPDLs, compared with CS and Ctl (p <
0.05). Angiogenesis is critical for hard tissue regeneration, and angiogenic induction
by bone substitutes itself is a simple and efficacious strategy for hard tissue
regeneration [14,51]. Angiogenesis during new bone formation has been examined in
in vivo and in vitro experiments with calcium silicate-based materials [17,52].
However, the mechanism by which Mg ions promotes cell activity, including initial
cell proliferation and differentiation, remains unclear. In other words, it is unclear
which factors from Mg contribute to its stimulation of cementogenesis and
angiogenesis. Among them, Mg5 and Mg10 released more Mg ions and greater
stimulatory effects on angiogenesis induction than CS. These results imply a
relationship between the induction of angiogenesis and Mg concentration [53].
19
4. Conclusions
The results of the present study of newly developed Mg-calcium silicate
cements by sol-gel that they are not only fast-setting and have good bioactivity, but
also degrade in SBF. The ionic products of these materials’ dissolution stimulate cell
proliferation, odontogenesis differentiation and angiogenesis. The degree of
promotion varies depending on the Mg ion concentrations in the culture medium. In
addition, Mg-CS also stimulates the proliferation of hPDLs in vitro and actively
promotes the secretion of odontogenic (DSPP and DMP-1) and angiogenic (vWF and
ang-1) proteins. The in vitro findings provide the essential theoretical basis for further
in vivo study of Mg-CS cements used as biomaterials in bone substitutes and bone
regeneration applications.
20
Acknowledgements
The authors acknowledge receipt of a grant from the Ministry of Science and
Technology grants (MOST 104-2314-B-039-004) of Taiwan. The authors declare that
they have no conflicts of interest.
21
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