the mechanical and biological properties of an injectable calcium phosphate cement-fibrin glue...

The Mechanical and Biological Properties of an InjectableCalcium Phosphate Cement-Fibrin Glue Composite forBone Regeneration

Geng Cui,1 Jie Li,2 Wei Lei,3 Long Bi,3 Peifu Tang,1 Yutian Liang,1 Sheng Tao,1 Yan Wang1

1 Institute of Orthopedics, General Hospital of PLA, Beijing 100853, People’s Republic of China

2 Department of Gynecology and Obstetrics, General Hospital of PLA, Beijing 100853, People’s Republic of China

3 Institute of Orthopaedics and Traumatology, Xijing Hospital, The Fourth Military Medical University, Xi’an 710032,People’s Republic of China

Received 24 March 2009; revised 2 July 2009; accepted 1 August 2009Published online 10 November 2009 in Wiley InterScience (www.interscience.wiley.com). DOI: 10.1002/jbm.b.31525

Abstract: Calcium phosphate cement (CPC) that can be injected to form a scaffold in situ

has promise for the repair of bone defects. However, its low-strength limits the CPC to non-

stress-bearing repairs. Fibrin glue (FG) with good sticking property and biocompatibility is

possible used to reinforce the CPC. The objective of this study was to investigate the effects of

FG on the mechanical and biological properties of CPC in an injectable CPC-FG composite.

The initial setting time of this CPC-FG was delayed compared with the CPC control at

different powder/liquid (P/L) mass ratio (p > 0.05). At a P/L of 5, the strength was (38.41 64.32) MPa for the CPC-FG, much higher than (27.42 6 2.85) MPa for the CPC alone (p <0.05). SEM showed bone marrow stromal cells (BMSCs) with healthy spreading and anchored

on the CPC-FG composite. After 14 days, the alkaline phosphatase (ALP) activity was (538 633) for the BMSCs on the CPC-FG and (517 6 27) for the BMSCs on the CPC alone. Both

ALPs were higher than the baseline ALP (93 6 10) for the undifferentiated BMSCs (p < 0.05).

The results demonstrate that this stronger CPC-FG scaffold may be useful for stem cell-based

bone regeneration in moderate load-bearing orthopedic applications. ' 2009 Wiley Periodicals,

Inc. J Biomed Mater Res Part B: Appl Biomater 92B: 377–385, 2010

Keywords: calcium phosphate(s); fibrin; biomechanics; cell proliferation; cell differentiation

INTRODUCTION

Treatments of osseous defects, delayed healing, and nonun-

ion, for example, from trauma, infection or tumor, remain

major problems in orthopedic surgery.1–3 Autografts and

allografts represent the current strategies for surgical inter-

vention and subsequent bone repair, but each possesses

limitations, such as donor-site morbidity with the use of

autografts and the risk of disease transmission with the use

of allografts.2–4 Tissue engineering is a promising approach

for the repair and regeneration of defective or damaged tis-

sues and organs. The basic principles involved in tissue-

engineered bone are the isolation and cultivation of cells,

the use of osteoinducing substances or growth factors, and

the placement of cells within suitable matrices or on scaf-

folds to support their growth.5 Two different types of scaf-

folds are used in tissue engineering for growth factor

delivery and cell adhesion, namely scaffolds made from

solid, prefabricated materials and scaffolds made from

materials that are injectable and hardened in situ.6

Injectable materials hold promise for tissue engineering

applications because they offer some advantages over solid

materials for certain indications. Injectable materials elimi-

nate the need for surgical intervention for their delivery,

and the minimally invasive procedure of injection reduces

discomfort and complications for the patient.7 Moreover,

injectable materials are able to take the shape of the cavity

into which they are placed and can thus fill irregular

defects. Also, problems with cell adhesion and growth fac-

tor delivery are overcome because, under the proper condi-

tions, the materials can be easily incorporated into the

solution by mixing prior to the injection.7,8 Candidates for

injectable materials include natural polymers, synthetic

polymers, and inorganic materials.6 Because of their simi-

larity to human bone, inorganic materials such as calcium

Correspondence to: Y. Wang (e-mail: [email protected])

' 2009 Wiley Periodicals, Inc.

377

phosphate ceramics and cements were, and still are, popu-

lar implant materials for diverse clinical applications.9–11

Different types of Ca-P ceramics are available and can be

classified as hydroxyapatite (HA), beta-tricalcium phos-

phate (b-TCP), or biphasic calcium phosphate (BCP).9

However, because they are made from sintered Ca-P

ceramics, they cannot be injected and must be shaped to

the defect site during operation, which makes the process

very labor intensive. A less than optimal tissue-to-implant

contact is often obtained. Because of the low-processing

temperature, calcium phosphate cements (CPCs), do not

comply with the original definition of calcium phosphate

ceramics, although the powder phase predominantly con-

tains TTCP with minor amounts of a-tricalcium phosphate

and HA.12,13 The main difference between cements and

ceramics is the injectability and in situ hardening of the

cement, which makes it easy to handle from a clinical point

of view.13,14 CPCs are obtained by mixing one or several

reactive calcium phosphate powders with an aqueous solu-

tion to form a paste that hardens within a restricted period

of time (e.g., 15 min). Two main approaches can be used

to make a CPC via an acid–base reaction or via a conver-

sion reaction of a metastable compound, either a-tricalcium

phosphate or a so-called amorphous calcium phosphate.13

However, due to its brittleness and weakness, the use of

CPC was limited to the reconstruction of non-stress-bearing

bone.15–17 Other injectable materials like chitosan, gelatin

and collagen are introduced to CPC to improve its mechan-

ical properties.18–20 However, studies have seldom been

done on the composite of CPC incorporated with fibrin

glue (FG) and its properties.21 FG is another type of inject-

able biomaterial approved by the Chinese State Food and

Drug Administration (SFDA) and is widely used in surgery

for repairing tissue damage or blood vessel injury.22–24 The

formation of FG involves the formation of a covalent iso-

peptide bridge between Gln and Lys residues by the enzy-

matic action of the transglutaminase factor XIIIa.24 FG

contains high concentrations of fibrinogen, which can pro-

duce a dense fibrin clot with sufficient adhesive strength to

maintain a required configuration for tissue adhesion and

cell ingrowth.25,26 Both CPC and FG have shown great

properties for the proliferation and differentiation of cells;

thus, the incorporation of FG was expected to increase the

strength without decreasing the biocompatibility of CPC.

We presented an injectable composite by combined FG

with CPC for the repair of bone defects. The objectives of

this study were to investigate the mechanical and biological

properties of this CPC-FG composite. The hypotheses are

as follows: (1) FG can increase the strength of this CPC-

FG composite compared with CPC, with little influence on

its setting process; (2) bone marrow stromal cells (BMSCs)

derived from rabbit bone marrow will attach to this CPC-

FG scaffold with a high-proliferation rate; and (3) the

CPC-FG scaffold will support the osteogenic differentiation

of BMSCs, yielding an elevated expression of alkaline

phosphatase (ALP) and collagen I (Col I).

MATERIALS AND METHODS

Preparation of the CPC-FG Composite

Preparation of the CPC powder was done according to

Ref. 18. Briefly, TTCP was synthesized from a solid-state

reaction between equimolar amounts of DCPA and CaCO3,

which were mixed and heated at 15008C for 6 h in a fur-

nace. The heated mixture was quenched to room tempera-

ture, ground in a ball mill, and sieved to obtain TTCP

particles with sizes ranging from �1 to 80 lm, with a me-

dian of 17 lm. DCPA was ground for 24 h to obtain parti-

cle sizes ranging from about 0.4 to 3 lm, with a median of

1 lm. The TTCP and DCPA powders were then mixed in

a blender at a molar ratio of 1:1 to form the CPC powder.

A commercially available SFDA-approved FG (Hualan

Biological Engineering, China) was used. The CPC powder

and FG solution were mixed at three different powder/liq-

uid mass ratios (P/L) (g/mL): 5/1, 3, and 1 in sterile injec-

tion syringes. Each CPC-FG paste was immediately

injected into a circular mold of 10-mm diameter and 2-mm

thickness to make disks for the setting time and cell studies

and a circular mold of 10-mm diameter and 20-mm high to

make cylinders for mechanical testing. Each specimen was

set in a humidor with 100% relative humidity at 378C for

4 h and then demolded and immersed in distilled water

378C for 20 h. The SFDA-approved CPC (Shanghai

Rebone, China) made using the same CPC powder but

using water as the liquid without the FG, was also fabri-

cated to serve as a control.

Setting Time with Gillmore Needles

Gillmore needles standard27 was used to compare the set-

ting time of the CPC-FG with that of the CPC (n 5 5).

Briefly, after the CPC-FG and CPC of different P/L ratios

were made in moulds (according to M&M 2.1), needles

were used to scrub the paste gently according to the refer-

ence. The initial setting time is defined as the time when a

0.3 MPa static pressure does not leave a visible print on

the surface of the cement. The final setting time is the cor-

responding time for a static pressure of 5 MPa. Each speci-

men was set with 100% relative humidity at 378C.

Mechanical Testing

The entire specimens were cylinders of 10-mm diameter

and 20-mm height (n 5 5). The mechanical testing was

performed after the specimens were left at room tempera-

ture for 24 h. The mechanical properties of the specimens

were analyzed in a materials testing machine (MTS-858,

MTS System, USA) by compression in the cranial-caudal

direction at a deformation rate of 5 mm/min until failure.

The compressive strength was calculated by S 5 Fmax/A,where Fmax is the maximum load on the load-deformation

curve and A is the cross-sectional area of each specimen.

The elastic modulus was calculated by E 5 S/(L/H), where

378 GENG ET AL.

Journal of Biomedical Materials Research Part B: Applied Biomaterials

L is the load deformation value and H is the initial height

of the specimen.

BMSCs Culture on CPC and CPC-FG

The BMSCs were prepared as described previously.28

Briefly, BMSCs were obtained from the os longum of

female white New Zealand rabbits aged between 8 months

and 1 year. All animal procedures used in this study were

approved by our institutional animal care committee. The

cells were cultured at 378C in a humidified atmosphere of

5% CO2 in flasks containing Dulbecco’s modified Eagle’s

medium (DMEM; Gibco, USA), 10% fetal bovine serum

(FBS; HyClone, USA), and 1% penicillin/streptomycin.

The medium was changed every third day. Cell subcultures

of third passages were used in the experiments. Osteogenic

differentiation was induced by culturing the BMSCs in an

osteogenic medium (OM) consisting of DMEM supple-

mented with 10% FBS, 1028 M DEX, 1023 M b-glycerolphosphate, and 50 mg/mL L-ascorbic acid. Prior to the cell

culture, the CPC control and the CPC-FG disks were steri-

lized by gamma radiation at 25 KGy. The disks were rinsed

in Dulbecco’s phosphate buffered saline. Upon reaching

90% confluence, cells adhesive to the flasks were detached

with 0.25 wt % trypsin (Type XI; Sigma Chemical, USA)

in 0.1M PBS. A total of 5 3 104 cells were diluted into

2 mL of OM and added to each well of a 24-well plate.

Each of the wells contained a disk of either CPC-FG or

CPC control.

SEM Observation

After being cultured for days 3 and 7, the specimens were

removed from the plates and gently washed with PBS three

times. The specimens were then fixed with 3% glutaralde-

hyde in PBS for 24 h at 48C. After being thoroughly

washed with PBS, the specimens were dehydrated sequen-

tially in 30, 50, 70, 80, 90, 95, and 100% ethanol. The

specimens were dehydrated twice in each ethanol concen-

tration for 15 min each time. The specimens were freeze

dried, sputter coated with gold, and examined under a SEM

(S-3000N, Japan) operated at 20 kV.

BMSCs Proliferation Assay

The number of BMSCs in the scaffold was determined

after 1, 7, and 14 days of triplicate culturing. In brief, at

these predetermined times, five specimens of each group

were gently washed with PBS to remove unattached cells.

The adherent cells were removed from the specimens by

incubation for 5 min in 2 mL 0.25 wt % trypsin in 0.1MPBS. Then, the suspension was collected and centrifuged at

300 g for 10 min. The sediment cells were resuspended in

fresh medium. One drop of this cell suspension was com-

bined with an equal amount of 0.4% (w/v) trypan blue

(Sigma Chemical, USA) and counted using a hemocytome-

ter under phase-contrast light microscopy.

Type-I Collagen (Col I) Immunohistochemistry Staining

The expression of Col I in the BMSCs was determined at

14 days by the Avidin-biotin peroxidase complex (ABC)

method. Briefly, the BMSCs on each specimen were

detached with 0.25 wt % trypsin (in 0.1M PBS) at 378C for

10 min. Then the specimens were extracted, and the cells

were centrifuged. The deposit of cells were resuspended in

PBS containing 10% fetal calf serum and the suspended cells

were incubated on 10 mm 3 10 mm sterile glass slides in

OM until adherence. Then the OM was removed. After speci-

mens were fixed with acetone at 48C for 10 min, they were

incubated in a moist chamber at 308C for 2 h with a primary

antibody, monoclonal rabbit anti-rat Col I diluted in PBS

with 1% BSA to 1:800. A horseradish peroxidase-3,30-diami-

nobenzidine tetrahydrochloride (HRP-DAB) system kit

(R&D System, USA) was used, including a secondary anti-

body, ABC reagent, and DAB system. The cells were incu-

bated with biotinylated secondary antibodies (rat anti-rabbit)

at room temperature for 30 min and incubated with the ABC/

HRP for 30 min. The results were observed by light micros-

copy at high magnification. BMSCs cultured in OM without

any specimen were incubated on the sterile glass slides and

regarded as positive control. On the other hand, BMSCs cul-

tured without OM and specimen were incubated on the ster-

ile glass slides, and were regarded as negative control.

ALP Measurement

The ALP activity was determined at 7 and 14 days of cul-

turing according to Ref. 18. Briefly, the medium was

removed and at the predetermined time point, five speci-

mens of each group were transferred to a new well plate.

Then, 0.5 mL of Triton X-100 was added to each well. A

cell scraper was used to remove the BMSCs from the disk

surface. The disk and 0.5 mL of cell lysate were placed in

a 1.5 mL centrifuge tube. The samples were then processed

through two freeze–thaw cycles (2808C and room tempera-

ture, 45-min each) to rupture the cell membranes and

extract the proteins and DNA from the cells. A p-nitro-phenyl phosphate (pNPP) liquid substrate system (Nanjing

Jiancheng, China) was used to analyze the ALP concentra-

tion from the cells on each disk. Five milliliters of each

cell lysate solution was added to 195 lL of pNPP substrate

and incubated in the dark at room temperature for 1 min.

The absorbance was read using a plate reader (Molecular

Devices, USA) at 405 nm and normalized to the PicoGreen

assay.29 The DNA was quantified using the PicoGreen Kit

(Invitrogen, USA) following standard protocols. Briefly,

100 lL of each cell lysate solution was added to 100 lL of

PicoGreen reagent and incubated in the dark at room tem-

perature for 5 min. The absorbance was read at an excita-

tion/emission of 480–520 nm on the plate reader.

Statistical Analysis

Statistical analyses were performed using SPSS software,

version 12.0 (SPSS, USA). The data were presented as the

379CPC-FG COMPOSITE FOR BONE REGENERATION


mean 6 SD, and levels were compared by the nonparamet-

ric Mann–Whitney U test or Student’s t-test. P-values less

than 0.05 were considered significant.

RESULTS

As shown in Figure 1, under identical conditions (378C,100% relative humidity), the setting time was delayed with

a decrease in P/L in the CPC-FG specimens and CPC con-

trols. The initial setting time of the CPC-FG specimens

(they were 20.41 6 1.21, 15.37 6 1.10, and 8.33 6

0.62 min at a P/L of 1, 3, and 5, respectively) was longer

than that of the CPC controls (they were 15.09 6 1.49,

10.97 6 1.52, and 5.22 6 0.16 min at a P/L of 1, 3, and 5,

respectively) (p\ 0.05), but there was no significant differ-

ence in the final setting time between the CPC-FG and the

CPC control (p[ 0.05).

The compressive strength and elastic modulus (mean 6SD; n 5 5) are plotted in Figure 2. The compressive

strength and elastic modulus declined with a decrease in

P/L in the CPC-FG specimens and CPC controls. The com-

pressive strength of the CPC-FG specimens (they were

Figure 1. The initial setting time of the CPC-FG specimens at different powder/liquid (P/L) mass ra-tio was longer than that of the CPC controls (*p \ 0.05), but there was no significant difference in

the final setting time between the CPC-FG and the CPC control (#p [ 0.05); (mean 6 SD, n 5 5).

[Color figure can be viewed in the online issue, which is available at www.interscience.wiley.com.]

Figure 2. The compression strength and elastic modulus declined with a decrease in P/L in the

CPC-FG specimens and CPC controls. The compression strength of the CPC-FG specimens was

higher than that of the CPC control at different P/L ratio (*p\ 0.05). The elastic moduli of the CPC-

FG scaffolds at a P/L of 1 and 3 were obviously higher than that of the CPC controls (*p\ 0.05). Ata P/L of 5, the modulus was not significantly different between the CPC-FG and CPC control (#p[0.05); (mean 6 SD, n 5 5). [Color figure can be viewed in the online issue, which is available at

www.interscience.wiley.com.]

380 GENG ET AL.


10.15 6 1.59, 26.93 6 3.26, and 38.41 6 4.32 MPa at a

P/L of 1, 3, and 5, respectively) was higher than that of the

CPC control (they were 7.00 6 1.33, 19.32 6 2.40, and

27.42 6 2.85 MPa at a P/L of 1, 3, and 5, respectively) (p\ 0.05). The elastic moduli of the CPC-FG scaffolds at a

P/L of 1, 3, and 5 (2.16 6 0.16, 2.73 6 0.24, and 3.20 60.16 GPa) were higher than that of the CPC controls (1.79

6 0.21, 2.28 6 0.14, and 3.03 6 0.14 GPa) (p\ 0.05).

The BMSCs showed good proliferation on both scaf-

folds, as observed by SEM (Figure 3). After 3 days, cells

adhered tightly to the CPC-FG and showed a uniform dis-

tribution. Their cytoplasmic extensions formed a meshwork

on the surface of the scaffold. The formation of the cell–

cell junctions indicated that good communication was

formed between the cells [Figure 3(b)]. After 7 days, the

cells were confluent on the surface of the CPC-FG scaffold.

The extracellular matrix secreted by the cells surrounded the

cells and the surface of the scaffold, which proved that

the cells had good conditions and the scaffold had good

biocompatibility [Figure 3(c,d)].

The proliferation of the BMSCs was measured after 1,

7, and 14 days by cell counting. From day 1 to 14, the

number of cells on the CPC-FG and the CPC control

increased (Figure 4). The number of cells on both of the

Figure 3. SEM photomicrographs of CPC-FG scaffold and CPC-FG with BMSCs at different time.The surface of CPC-FG showed rough with some pores (a). After 3 days of culture, cells adhered

tightly to the CPC-FG and showed a uniform distribution. Their cytoplasmic extensions formed a

meshwork on the surface of the scaffold. The formation of the cell–cell junctions indicated that

good communication was formed between the cells (b, arrows). After 7 days, the cells were conflu-ent on the surface of the CPC-FG scaffold (c). The extracellular matrix secreted by the cells sur-

rounded the cells and the surface of the scaffold (d, arrows). The surface of the CPC control was

irregular with some pores (e) and plenty of BMSCs extended on the surface of the CPC at the 7th

day (f, arrows).



materials did not differ (p \ 0.05). These results demon-

strate that the CPC-FG scaffold supported BMSC prolifera-

tion equal to that of the SFDA-approved CPC control.

Immunohistochemically (Figure 5), the brown area

stained by DAB was a positive result, which indicated the

expression of collagen type I. After 14 days, the high

expression of collagen type I was localized within BMSCs

under the OM [Figure 5(a)] and low expression of collagen

type I was localized within BMSCs without the OM

[Figure 5(b)].

The ALP, normalized to DNA concentration, is plotted

in Figure 6. Under an OM, the ALP expression on each

scaffold increased substantially from day 7 to 14. There

was no significant difference on days 7 and 14 between the

CPC-FG [244 6 23 and 538 6 33 (pNPP mM/min)/(mg

DNA) on days 7 and 14, respectively)] and the CPC con-

trol [230 6 17 and 517 6 27 (pNPP mM/min)/(mg DNA)

on days 7 and 14, respectively] (p [ 0.05). Both values

were much higher than those of the baseline ALP for

BMSCs in control medium without the OM (p\ 0.05).

DISCUSSION

A major disadvantage of conventional materials in the field

of bone tissue engineering is that they exist in a hardened

form, requiring the surgeon to drill the surgical site around

the implant or to carve the graft to the desired shape.29

This can lead to increases in bone loss, trauma, and surgi-

cal time.18,30 To realize minimally invasive injury and

facilitate operative procedures, injectable biomaterials are

widely researched.31–33 The requirements of injectable scaf-

folds are numerous, but most important are that they must

be biocompatible and mechanically stable. Because of their

similar inorganic components to human bone, CPC have

good biocompatibility and hold promise for diverse clinical

applications. CPC possesses excellent osteoconductivity

and bone replacement capability and is highly promising

for use in a number of restorative dental and craniofacial

procedures. The FDA approved it for craniofacial indica-

tions a decade ago. However, when CPC was used in peri-

odontal bone repair, tooth mobility resulted in early

fracture and eventual exfoliation of the rigid and brittle

implants. CPC is a typical brittle ceramic and can fracture

catastrophically at a small deformation strain of about

0.2%. Therefore, it is desirable to have CPC in a nonrigid

form that can sustain large deformation strains without

fracture.34 The additions of other injectable materials like

Figure 4. From day 1 to 14, the number of cells on the CPC-FG

and the CPC control increased. The number of cells on both of the

materials did not differ (#p [ 0.05); (mean 6 SD, n 5 5). [Color fig-ure can be viewed in the online issue, which is available at www.

interscience.wiley.com.]

Figure 5. After being stained by DAB at the 14th day, amounts of brown particles could be found

in the cytoplasm of BMSCs under the OM (a, arrows) and seldom brown particles could be found

within BMSCs without the OM (b). [Color figure can be viewed in the online issue, which is available

at www.interscience.wiley.com.]

382 GENG ET AL.


chitosan, gelatin, and collagen to CPC have resulted in

composites of CPC with high-mechanical properties.18–20

A commercially available SFDA-approved FG was com-

bined with CPC to improve the mechanical properties of

the CPC in this study. We found that the CPC-FG compos-

ite possessed a much higher compressive strength than the

scaffold made only from the SFDA-approved CPC control.

The incorporation of FG proved to significantly increase

the compressive strength of the CPC. FG is an elastomeric,

biocompatible, and natural biomaterial with sufficient adhe-

sive strength. The commercially available FG includes two

main components: one part is fibrinogen and the other part

is thrombin. The former contains fibrinogen, fibrin-stabiliz-

ing factor XIII, and aprotinin, whereas the latter contains

thrombin and calcium chloride. When the two parts are

mixed, the fibrinogen is hydrolyzed by thrombin and

changed into a fibrin monomer. The fibrin monomer can

polymerize into soluble fibrin by the hydrogen bond and

electrostatic attraction. In the presence of Ca21, factor XIII

is activated by thrombin, and the noncovalent bond

between the fibrin monomer changes into a covalent bond

by transamination. The result is a stable, high-strength

fibrin clot (i.e., FG is formed), which is not easily degraded

by fibrinolysin and has good coherence. When the CPC

and the two parts of FG were mixed, the CPC powder

could stick together with the fibrin of the FG to form a

three-dimensional porous network structure in the compos-

ite. The network structure could be observed by SEM

[Figure 3(a)]. The traditional CPC (without FG) relied

solely on the apatite crystals to interlock to provide

strength. In contrast, the CPC-FG composite had not only

interlocking crystals, but also additional FG to bind the ap-

atite crystals together. This interlacing fibrin clot helped to

increase the strength and flexibility of the CPC-FG com-

posite.

Setting time is another index used to evaluate injectable

scaffolds.18,35,36 Fast setting enables the graft to attain a

significant strength and geometric integrity within a short

period of time postoperatively. The addition of other com-

ponents, such as growth factors and other materials, to the

liquid phase could delay the CPC setting.9 In our study,

we found that the initial setting time of the CPC-FG com-

posite was obviously delayed, but the final setting exhibited

little change. The reason may be that the incorporation of

the FG had an effect on the interactions between the CPC

powders. After 3–5 min, when the stable fibrin clot was

formed, the effect of the FG was decreased, the interaction

between the CPC powders was recovered and the coher-

ence of the FG could help in the setting procedure of CPC

powders. When carrying out the setting procedure, we also

found that the injectable property of this composite could

be controlled by the ratio of P/L. At a P/L of 1, the CPC-

FG paste flowed easily and could be injected through a

small 18-gauge needle. The paste at a P/L of 3 could be

injected using a 12-gauge needle. The paste was injectable

even at a P/L of 5, and the strength was higher than that at

lower P/L ratios.

Ideally, the scaffold should be bioactive so that it can

promote and guide cell proliferation, differentiation, and

tissue growth. The CPC was regarded as holding good bio-

compatibility.2,18,33,35 This study showed that the CPC-FG

scaffold could support BMSC attachment and proliferation,

matching the SFDA-approved CPC. After 3 days, the

BMSCs were able to adhere and spread on the CPC-FG.

The BMSCs attained a normal morphology on the scaffold

and formed cell–cell junctions. After 7 days, the BMSCs

greatly proliferated, increasing the cell number by an order

of magnitude. Hence, the CPC-FG scaffold had higher

strength, without compromising the colonization and prolif-

eration of BMSCs, compared with the CPC control. Both

ALP and Type-I collagen are expressed by BMSCs during

osteogenesis and are well-defined markers for their differ-

entiation.37,38 In this research, the ALP of the BMSCs on

both the CPC and the CPC-FG scaffolds in the OM showed

an increase of 5- to 6-fold to an order of magnitude, com-

pared with the ALP of BMSCs without the OM. We also

observed a high expression of Type-I collagen from the

BMSCs cultured on the CPC-FG and CPC scaffolds in the

OM. Hence, the addition of FG to CPC, which strength-

ened the CPC, did not compromise the ALP and Type-I

collagen secretions of the BMSCs. Therefore, the BMSCs

attaching on the CPC and CPC-FG scaffolds were success-

fully differentiated down the osteogenic lineage and

expressed elevated levels of osteogenic markers in vitro.Although good results have been obtained during

in vitro research, further studies are necessary. First, the effectof the FG in the CPC scaffold on cell behavior needs to be

understood. In addition, studies are needed to incorporate

the stem cells into the CPC paste for injection delivery,

Figure 6. ALP activity, normalized to DNA concentration, with units

of (pNPP mM/min)/(mg DNA). Under an OM, the ALP expression on

each scaffold increased substantially from day 7 to 14. There wasno significant difference on days 7 and 14 between the CPC-FG

and the CPC control (#p [ 0.05). Both values were much higher

than those of the baseline ALP for BMSCs in control medium with-

out the OM (*p \ 0.05); (mean 6 SD, n 5 5). [Color figure can beviewed in the online issue, which is available at www.interscience.

wiley.com.]



instead of seeding the cells on the surface of a prehardened

CPC. Furthermore, because the in vitro cell culture is a

closed static system and differs from the dynamic circula-

tion in vivo, animal studies are needed to investigate the

bone regeneration efficacy of the CPC-FG-stem cell

construct.

CONCLUSIONS

The CPC-FG scaffold possessed compressive strength

higher than the CPC control. The CPC-FG composite sup-

ported BMSC attachment and proliferation. Although the

initial setting time of the CPC-FG composite was obviously

delayed compared with the CPC control, the final setting

time had little change. The expression of ALP and Type-I

collagen of the BMSCs in this CPC-FG in the OM was

also increased compared with the BMSCs in control me-

dium without osteogenic factors. Therefore, the addition of

FG to CPC, which strengthened the CPC, did not compro-

mise the proliferation and differentiation of BMSCs.

Hence, the CPC-FG scaffold may have potential for BMSC

delivery and bone regeneration in moderate stress-bearing

orthopedic applications.

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the mechanical and biological properties of an injectable calcium phosphate cement-fibrin glue...

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