Acta Biomaterialia 6 (2010) 845–855

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Acta Biomaterialia

journal homepage: www.elsevier .com/locate /actabiomat

Reactive calcium-phosphate-containing poly(ester-co-ether) methacrylatebone adhesives: Chemical, mechanical and biological considerations

Xin Zhao, Irwin Olsen, Haoying Li, Kris Gellynck, Paul G. Buxton, Jonathan C. Knowles, Vehid Salih,Anne M. Young *

Division of Biomaterials and Tissue Engineering, UCL Eastman Dental Institute, 256 Gray’s Inn Road, London WC1X 8LD, UK

a r t i c l e i n f o a b s t r a c t

Article history:Received 9 June 2009Received in revised form 24 September 2009Accepted 28 September 2009Available online 1 October 2009

Keywords:Poly(propylene glycol-co-lactide)dimethacrylateCalcium phosphateMechanical propertiesBiocompatibilityBone repair

1742-7061/$ - see front matter � 2009 Acta Materialdoi:10.1016/j.actbio.2009.09.020

* Corresponding author. Tel.: +44 20 7915 2353; faE-mail address: a.young@eastman.ucl.ac.uk (A.M.

A poly(propylene glycol-co-lactide) dimethacrylate adhesive with monocalcium phosphate monohydrate(MCPM)/b-tricalcium phosphate (b-TCP) fillers in various levels has been investigated. Water sorption bythe photo-polymerized materials catalyzed varying filler conversion to dicalcium phosphate (DCP). Poly-mer modulus was found to be enhanced upon raising total calcium phosphate content. With greater DCPlevels, faster release of phosphate and calcium ions and improved buffering of polymer degradation prod-ucts were observed. This could reduce the likelihood of pH-catalyzed bulk degradation and localized acidproduction and thereby may prevent adverse biological responses. Bone-like MG-63 cells were found toattach, spread and have normal morphology on both the polymer and composite surfaces. Moreover,composites implanted into chick embryo femurs became closely apposed to the host tissue and didnot appear to induce adverse immunological reaction. The above results suggest that the new compositematerials hold promise as clinical effective bone adhesives.

� 2009 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.

1. Introduction

With the rapidly increasing human survivability and life expec-tancy, there is a growing requirement for clinically effective adhe-sives and fillers to replace/repair damaged bone in orthopedic andoral surgery. Ideally, such materials should be injectable, fast set-table, degradable and biocompatible, and have a stiffness compara-ble with the surrounding tissue. Injectability enables the materialto accommodate to defects of different shapes, while fast settabil-ity and stiffness provide instant mechanical support to the sur-rounding tissue. Degradability enables new bone to progressivelyintegrate and ultimately replace the degrading material, eliminat-ing the need for revision surgery. With regard to biocompatibility,a number of parameters have previously been used to delineatethis property. In vitro measurements include cell attachment andproliferation [1,2] on the implant surface and in vivo vasculariza-tion of bone biomaterial implants [3], bone mineralization andintegration with the implanted material [4]. In the present study,biocompatibility is considered to be the adhesion of target cells[1,2] and tissues [5] to the material surface in vitro and in vivo.Such adhesive interactions have been shown to play a critical rolein cell survival, proliferation, differentiation, matrix mineralizationand, therefore, new bone formation [6]. The release of phosphate

ia Inc. Published by Elsevier Ltd. A

x: +44 20 7915 1227.Young).

and calcium ions is also considered to be biologically beneficialas it stimulates osseous tissue development [7].

Synthetic injectable degradable polymers such as poly(ether-co-esters) meet some of the above requirements and could thusbe of potential value in tissue regeneration. Such fluid materials,when modified with acrylate or methacrylate (MA) end groups,can be readily converted into solids after injection [8]. The cross-linkable acrylate or MA end-capped poly(ether-co-esters) pro-duced thus far contain primarily hydrophilic poly(ethylene glycol)(PEG) [9]. After cross-linking and subsequent placement in aque-ous fluids, these solidified materials, because of their high hydro-philicity, can degrade ‘‘catastrophically” in a manner similar tosome polylactides [10,11]. This may lead to a sudden loss of func-tional support for the damaged tissue and sometimes adverse hostreaction due to localized acid production [12,13]. In addition, thehighly hydrophilic characteristics of such PEG-based materials of-ten do not support optimal attachment and spreading of anchor-age-dependent cells [14], which may affect integration of thebiomaterial with the host tissue.

Poly(propylene glycol) (PPG)-containing polymers, for example,fluid poly(lactide-co-PPG-co-lactide) end-capped with MA groups,have therefore been developed since PPG has a similar molecularstructure to PEG but more methyl groups which endow a higherdegree of hydrophobicity [15]. Such materials are not only inject-able and rapid setting, but also likely to degrade and release acidmore slowly in water because of their higher hydrophobicitycompared with their PEG counterparts. These materials, however,

ll rights reserved.

846 X. Zhao et al. / Acta Biomaterialia 6 (2010) 845–855

are of relatively low modulus and their biological reactivity withbone-related cells and tissue is not yet known [15].

In order to improve the mechanical properties of polymericmaterials, biocompatible hydroxyapatite (HA) and b-tricalciumphosphate (b-TCP) particles have previously been incorporated as‘‘fillers” [16,17]. Although these fillers may increase material mod-ulus, they have low aqueous solubility at physiological pH (7.2–7.4) and thus limited pH compensation ability [18–20]. Other stud-ies have shown, however, that it is possible to incorporate moresoluble calcium phosphates into polymers as fillers [21]. For exam-ple, dicalcium phosphates (brushite or monetite) have higheraqueous solubility and may therefore be able to provide higherphosphate and calcium concentrations in solution with bufferingcapacity to mitigate the potentially adverse effects of excess acidproduction upon polyester degradation. Additionally these ionsare essential for bone re-mineralization [22–24].

In the present study, various amounts of monocalcium phos-phate monohydrate (MCPM) and b-TCP (reactants in brushite-forming bone cements) were incorporated into a fluid poly(lac-tide-co-PPG-co-lactide) dimethacrylate monomer as fillers in orderto fabricate composite materials (polymer + filler). These couldhave improved mechanical properties and reduced acid productioncompared with the unfilled polymer. It was hypothesized thatwater sorption by the cross-linked polymer could promote conver-sion of these fillers to brushite and/or monetite. Thus the first aimof the present study was to characterize any induced changes in fil-ler chemistry upon water sorption via Raman mapping and X-raydiffraction (XRD). Additionally, water sorption and elastic modulus(by compression) of the polymer and composites before and afterwater immersion were also examined. Subsequently, total massloss, phosphate and calcium ion release and pH variation of thecomposite storage solution were measured to determine the re-mineralization and buffering capacities of the calcium phosphatefillers. Finally, biocompatibility of these composites was assessedin vitro and in vivo to determine whether these materials couldbe suitable as potential bone adhesives.

2. Materials and methods

2.1. Materials

b-TCP (95%), triethylamine (98%) and methacryloyl chloride(97%) were purchased from Fluka, Gillingham, UK. MCPM (95%),poly(propylene glycol) (PPG, average molecular weight =1000 g mol�1), camphorquinone (CQ, 97%), N,N-dimethyl-p-tolui-dine (DMPT, 99%) and hydroxyethyl methacrylate (HEMA, 99%)were from Sigma–Aldrich, Gillingham, UK. D,L-Lactide was pur-chased from Purac, Gorinchem, The Netherlands. Dulbecco’s mod-ified Eagle’s medium (DMEM), fetal bovine serum (FBS),

Table 1Water sorption, compressive modulus and in vitro biocompatibility of composites and po

Composites q (%) m (%) E0

F% (%) T/M Md (lm)

70 4 90 13 ± 1 4 ± 1 2770 4 30 12 ± 1 4 ± 1 2170 1 90 21 ± 2 12 ± 1 2170 1 30 20 ± 1 12 ± 1 2760 2 60 12 ± 1 6 ± 1 1550 4 90 9 ± 1 4 ± 1 1250 4 30 7 ± 1 4 ± 1 1550 1 90 12 ± 1 10 ± 1 1250 1 30 11 ± 1 10 ± 1 14Polymer 4 ± 0.5 3 ± 0.3 9

q, water sorption after 24 h of water immersion; m, total mass loss at 12 weeks; E0, initia�DE24, reduction in modulus after 24 h of water immersion; Ef, final wet compressive m

phosphate-buffered saline (PBS, Ca2+/Mg2+ free), penicillin, strepto-mycin and Trypsin–EDTA were all purchased from Invitrogen, Pais-ley, UK. MCPM particles were ground using a ball mill (RetschMixer Mill MM 301, Haan, Germany) at 30 Hz, shaking frequencyfor 4 min and sieved using sieves of 20, 38, 75 and 106 lm (Labo-ratory test sieve, Endecotts Ltd., London, UK) in a vibratory sieveshaker (Fritsch Analysette 3 Spartan, Obertein, Germany) for30 min to obtain particles of diameters: 20–38 (denoted as 30),38–75 (60) and 75–106 (90) lm.

2.2. Synthesis of monomer

The monomer was synthesized by a modification of the proce-dure previously described [15]. Briefly, PPG (0.1 mol) was reactedwith D,L-lactide (LA) (0.4 mol) at 150 �C in a nitrogen atmospherefor 6 h, using stannous octoate (0.05 wt.% of PPG) as a catalyst.The intermediate product was then end-capped at both ends withMA units using 0.4 mol triethylamine and 0.4 mol methacryloylchloride, added drop-wise in an ice bath. The final product waswashed with acetone and extracted using hexane, then washedwith HCl and NaHCO3 solution (0.1 mol l�1 for both) and distilledwater to remove impurities, e.g., triethylamine HCl. The molecularstructure of the synthesized monomer was confirmed using nucle-ar magnetic resonance (NMR) spectroscopy (600 MHz Varian UnityINOVA Spectrometer, Palo Alto, USA) as previously reported [15].

2.3. Fabrication of composite materials

The synthesized oligomeric monomer of poly(propylene glycol-co-lactide) dimethacrylate (90 wt.%) was mixed with CQ (1 wt.%),DMPT (1 wt.%) and HEMA (8 wt.%). Various amounts of MCPMand b-TCP were added to the monomer mixture and the resultingcomposite mixtures moulded using steel rings, with top and bot-tom surfaces covered by acetate sheets (3 M AF 4301, Manchester,UK). They were cured in a light box (Triad� 2000TM visible light curesystem, Dentsply Trubyte, Palo Alto, USA) using blue light exposure(100 mW cm�2) for 12 min, to produce solid discs (12 mm diame-ter, 2 mm thick, unless otherwise specified). Each test was per-formed in triplicate, and results were expressed as mean ±standard deviation (SD).

There were nine formulations of the composites, as shown inTable 1, which were based on two-level, three-factor factorial de-sign with one intermediate formulation. The three variables werefiller mass fraction (F%), molar ratio of b-TCP to MCPM (T/M) andMCPM particle size (Md). Each variable had a high and low level,e.g., 70 and 50 wt.% for F%. Factorial design was used to selectthe formulation with optimum mechanical properties and pHcompensation ability. Data analysis of the effects of the three vari-

lymer.

(MPa) E24 (MPa) �DE24 (%) Ef (MPa) Cv (%)

± 4 9 ± 1 67 ± 3 6 ± 1 88 ± 15± 4 10 ± 1 52 ± 7 8 ± 1 85 ± 13± 4 7 ± 1 67 ± 9 6 ± 1 81 ± 13± 3 8 ± 1 70 ± 8 7 ± 1 80 ± 16± 1 9 ± 1 40 ± 4 6 ± 1 85 ± 13± 2 8 ± 1 33 ± 4 5 ± 1 85 ± 15± 1 10 ± 1 33 ± 3 7 ± 1 85 ± 14± 2 7 ± 1 42 ± 5 5 ± 1 81 ± 12± 3 8 ± 1 43 ± 10 6 ± 1 80 ± 12± 1 3 ± 0.5 22 ± 4 2 ± 0.5 88 ± 13

l dry compressive modulus; E24, compressive modulus after 1 day water immersion;odulus; Cv, relative cell viability. Data = average ± SD.

X. Zhao et al. / Acta Biomaterialia 6 (2010) 845–855 847

ables on different material properties is detailed in Section 2.9below.

2.4. Chemical characterization

The light cured composite specimens were soaked statically in10 ml of deionized water (pH 7, adjusted by 0.01 mol l�1 NaOH)at 37 �C. Deionized water was used in order to be able to readilydetect the phosphate and calcium ions as well as pH change ofthe storage solution arising from possible material degradation(detailed in Section 2.7). Specimens after 24 h and 12 weeks ofwater immersion were collected and air-dried at room tempera-ture for chemical characterization by Raman and XRD. During the12-week storage period, the solution was changed weekly to avoidsaturation of degradation products (e.g., Ca2+ and PO3�

4 ) in the stor-age solution.

2.4.1. Raman analysisRaman analysis was carried out using a LabRam spectrometer

(Horiba Jobin Yvon, Middlesex, UK). Maps of an area of(100 � 100 lm) of the surface of the samples before and afterwater immersion for 24 h and 12 weeks were obtained at a stepsize of 5 lm, using a 633 nm laser, 50� objective and 1800 grating.The mean Raman spectra of corresponding Raman maps were ob-tained following background subtraction and normalization usingthe polymer peak at 1447 cm�1 via LabRam software. The peakintensities due to other components then provide an indicationof the level of that component relative to the polymer [25,26].Standard Raman spectra of the polymer, MCPM, b-TCP, brushiteand monetite were also recorded.

2.4.2. XRDXRD (Brüker D8 Advance Diffractometer, Karlsruhe, Germany)

was used to identify the possible presence of different phases ofcalcium phosphates in the specimens. XRD data were collectedfrom 10� to 100� 2h with a step size of 0.02� and a count time of18.9 s, using Ni filtered Cu Ka radiation. Identification of calciumphosphates was by comparison with standards from the Interna-tional Centre for Diffraction Data database using CrystallographicaSearch-Match software (Oxford Cryosystems, Oxford, UK).

2.5. Water sorption and mass loss

The initial water sorption (at 24 h) and final total mass loss (at12 weeks) of the polymer (used as control) and composite discswere determined after placement of the specimens in neutraldeionized water at 37 �C. The mass of the initial dry specimens(W0) and swollen (Ws) and dry (Wd) discs after 24 h and 12 weeksof water immersion were measured using an electronic balance(Mettler Teledo, Osaka, Japan). The water sorption, q (%), was de-fined as the mass of absorbed water as a percentage of the initialdry mass

q ¼WS �Wd

W0� 100 ð1Þ

The mass loss, m (%), was calculated using

m ¼W0 �Wd

W0� 100 ð2Þ

2.6. Compressive modulus

Specimens (5 mm diameter, 2 mm thick) were collected beforeand after placement in 2.5 ml of neutral deionized water at 37 �Cfor 1, 4 and 7 days and then weekly up to 10 weeks. At each timepoint, the storage solution was replaced with 2.5 ml of refresh

deionized water. The variation of polymer (used as control) andcomposite compressive modulus with time after placement inwater was characterized using a dynamic mechanical analyzer(DMA 7e, Perkin-Elmer Instruments, Bucks, UK). A linearly increas-ing quasi-static stress was applied to the specimen on a parallelplate incorporating a 10 mm probe. An increasing static force from10 to 8000 mN was applied to the sample at a constant rate of500 mN min�1. The resultant displacement in strain was plottedagainst the applied stress, and the modulus (E) was calculated asthe slope of the initial linear portion of the stress–strain curveusing Pyris TM version 5 software.

2.7. pH measurement and phosphate and calcium ion release

The composite specimens were immersed in neutral deionizedwater at 37 �C and the specimen storage solutions collected and re-freshed after 2, 4, 6 and 24 h, 2, 4 and 7 days and then weekly up to12 weeks. pH of the collected solutions were determined using apH meter (Jenway 3340, Essex, UK) and phosphate and calcium re-lease monitored using ion chromatography with Chromeleon�

software (Dionex, Surrey, UK). For pH measurement, the storagesolution of polymer was used as control. To measure PO3�

4 , anICS2500 system equipped with an AS50 autosampler and anEG50 eluent generator system was used. The mobile phase was a30 mM KOH solution with a flow rate of 1.5 ml min�1 and 10 minrunning time. An ASRS�-300 (4 mm) column mounted in an AS50thermal compartment and an ASRS suppressor was used. Ca2+

was measured on a Dionex ICS1000 system equipped with anAS50 autosampler, using 20 mM methylsulphonic acid as the mo-bile phase at a flow rate of 1 ml min�1 with a CSRS ULTRA (4 mm)column and a CAES suppressor. The running time was also 10 min.Calibration was obtained with standard solutions containing 1, 10,25, 50 and 100 ppm of PO3�

4 or Ca2+. An injection loop of 25 ll wasused.

2.8. Biocompatibility

In vitro biocompatibility was examined using the MG-63 osteo-sarcoma cell line, which has previously been widely employed asan in vitro test system for assessing the effects of many types ofbiomaterials [27,28]. The cells were cultured in polystyrene flasksin DMEM supplemented with 10% FBS, 50 IU ml�1 of penicillin and50 lg ml�1 of streptomycin at 37 �C in a humidified atmosphere of5% CO2 in air. Sub-confluent cells were passaged by using trypsin–EDTA (0.25% (w/v) trypsin, 1 mmol l�1 EDTA).

Polymer and composite films were fabricated by spreading0.03 g of the fluid unpolymerized materials onto a glass coverslip(13 mm diameter, VWR, Leicester, UK), covering by acetate sheet,curing in a light box for 2 min and sterilizing under UV light(254 nm) for 30 min. The films (attached to glass coverslips) weresubsequently placed into individual wells of 24-well tissue cultureplates and incubated in 1 ml of DMEM at 37 �C for 24 h. Followingmedium removal, the films were washed twice with 1 ml of freshmedium and inoculated with 1 ml aliquot of MG-63 cells (10,000cells ml�1) and incubated at 37 �C for 3 days. Following mediumaspiration, the cells were washed with PBS and fixed using 3% (v/v) glutaraldehyde in 0.1 M sodium cacodylate buffer for 24 h. Thespecimens were finally dehydrated in a graded series of ethanol(50, 70, 95 and 3� 100%) for 10 min each, immersed in hexameth-yldisilazane for 2–3 min, air-dried for 3 h, mounted and coatedwith gold/palladium. Cell attachment and morphology was ob-served by scanning electron microscopy (SEM, JEOL JSM-5410LV,Tokyo, Japan) at an accelerating voltage of 15 kV. Cell growthwas evaluated by direct counting of the number of cells on eachsubstrate as a percentage of the number of cells present on

848 X. Zhao et al. / Acta Biomaterialia 6 (2010) 845–855

ThermanoxTM plastic coverslips (used as control, NUNCTM, ThermoFisher Scientific, Loughborough, UK).

The in vivo biocompatibility of the set materials was evaluatedusing a chicken chorioallantoic membrane (CAM) model as previ-ously described [5,29]. Composite ‘‘films” with F = 5%, T/M = 1and Md = 30 lm were produced and sterilized as described above.After light curing and sterilization, the film was cut into strips(4 mm long, 1 mm wide) prior to implantation.

Femurs were isolated from fertilized eggs (J.K. Needle and Co.,Cheshunt, UK) after 14 days of incubation at 39 �C. Soft tissuewas removed from the femur and a small defect (4 mm long andup to the bone marrow cavity) was made manually with a tip ofa needle (25G, BD Microlance, Erembodegem, Belgium) in the mid-dle of the femur, the periosteum disrupted only at the defect site.The films were implanted sagittally and gently tapped with forcepsinto the prepared defect so that the surfaces of the film and femurwere level. The femur and film was then placed onto the CAM of a7-day-old host egg which had previously been ‘‘windowed” at day3 [30]. The windows were sealed with tape and the host eggs incu-bated at 39 �C for a further 7 days. Two non-implanted femurs(negative control), two femurs implanted with HA-spheres (posi-tive control) and four test femurs implanted with one compositestrip in each were placed in the host eggs (two femurs per egg).The HA beads were synthesized and produced according to Ref.[31]. HA was considered as positive control in this study due toits similar composition and structure to natural bone mineraland its biocompatibility and osteoconductivity [32]. After 7 days,the femurs were collected and fixed with 4% paraformaldehydefor 24 h, washed with PBS prior to dehydration in ascending con-centrations of ethanol, then xylene and finally embedded in wax.Transverse 8 lm sections were cut, placed on a glass slide, im-mersed in xylene to remove wax and rehydrated in descendingethanol concentrations and stained with toluidine blue.

a

02468

101214161820

750 950 1150 1350 1550Raman shift (cm-1)

Inte

nsity

(a.u

.) 12 week

24 h

0 h

C

0

1

2

3

4

5

6

7

8

750 950Rama

Inte

nsity

(a.u

.)

Fig. 1. Raman spectra of representative formulations of composite specimens (a) (F% = 7and after 24 h and 12 weeks of immersion in water. The spectra of the polymer, b-TCP,

To further determine any gross cytotoxicity of the compositeimplant, two composite strips in the absence of femurs wereplaced directly onto the CAM of a 7-day-old host egg (one stripper egg) and incubated at 39 �C for a further 7 days. As a control,poly(lactic-co-glycolic acid) (PLGA) microporous spheres were pre-pared according to Ref. [33] and used because of its biocompatibil-ity and wide clinical application [33,34].

2.9. Factorial analysis

Factorial analysis was used in this study to assess the effect ofthe three variables (F%, T/M and Md) on material properties (P inEq. (2)) including water sorption, compressive modulus and phos-phate and calcium release. An appropriate factorial expression inthis case would be [35,36]:

ln P ¼ hln Pi � a1 � a2 � a3 ð3Þ

where hln Pi is the average value of ln P for all eight possible formu-lation combinations excluding the intermediate formulation. ai

indicates the level of an effect a variable (e.g., F%) has on a materialproperty (e.g., water sorption). A positive ‘‘a value” means that theproperty increases with increasing variable and vice versa for neg-ative ai. Higher values of ai indicate greater effects of the variable ona property. ai for each term was determined using the followingequation:

ai ¼12

lnPH

PL

� �ð4Þ

where PH and PL are the geometric means of a property of all sam-ples with the variable, e.g., F%, at its high (H) (e.g., 70%) and low le-vel (L) (e.g., 50%), respectively. Values of ‘‘a” were determined intriplicate, and results expressed as means ± SD.

b

0

1

2

3

4

5

6

7

8

750 950 1150 1350 1550Raman shift (cm-1)

Inte

nsity

(a.u

.)

12 week

24 h

0 h

1150 1350 1550n shift (cm-1)

monetite

brushite

MCPM

β-TCP

polymer

0%, T/M = 4 and Md = 30 lm) and (b) (F% = 70%, T/M = 1 and Md = 30 lm) before (0 h)MCPM, brushite and monetite are shown in (c) as standards for comparison.

a

X. Zhao et al. / Acta Biomaterialia 6 (2010) 845–855 849

2.10. Statistics evaluation

The statistical significance of differences between in vitro bio-compatibility for each formulation was evaluated using one-wayANOVA. Since the relative cell viability values were homoge-neously distributed, the Bonferroni post hoc test was used. Datawere evaluated using SPSS 14.0 for Windows (SPSS, Inc., Chicago,IL, USA). The results were expressed as mean ± SD, and p values<0.05 were considered statistically significant.

b

Fig. 2. XRD spectra of representative formulations of composite specimens (a)(F% = 70%, T/M = 4 and Md = 30 lm) and (b) (F% = 70%, T/M = 1 and Md = 30 lm)before (0 h) and after 24 h and 12 weeks of placement in water. s b-TCP, h MCPM,d monetite, j brushite.

3. Results

The molecular structure of the monomer synthesized in thepresent study was comparable with previous reports [15], as con-firmed by NMR (data not shown). The efficiency of LA attachmentto PPG and MA end-capping was 97% and 95%, respectively. Thecomposites with F% = 50 and 60% were fluid whereas specimenswith F% = 70% were thicker and paste-like (did not spread whenplaced on a flat surface).

3.1. Chemical changes before and after water immersion

No characteristic ‘‘C@C” peak at 1640 cm�1 was observed in anyof the polymer or composite Raman spectra, consistent with fullmonomer conversion during light exposure (spectral region notshown). Inorganic Raman peaks were easier to distinguish in com-posites with higher F% and due to the relatively small areas exam-ined, more reproducible with smaller Md. The major factor affectingchanges in the average Raman spectra with time was T/M (see, forexample, Fig. 1). With excess b-TCP (particularly T/M = 4), its peaksat 945/970 and 1046 cm�1 dominated the average Raman spectra.Some MCPM could be observed in Raman maps of initial samplesand regions of brushite/monetite after 24 h (data not shown). Peaksdue to these components in the average spectra were, however, dif-ficult to detect due to the highly dominant b-TCP peaks.

Conversely, with equimolar b-TCP and MCPM (doublet at 903/915 cm�1), peaks in average Raman spectra and particles in Ramanmaps due to both components could readily be detected in drysamples. Following 24 h water immersion, the MCPM peaks van-ished. b-TCP peak height relative to that of the polymer also re-duced by more than 60%. In their place a shoulder at 980 cm�1

attributed to brushite and/or monetite could be observed. Theseresults were consistent with some of the MCPM dissolving butmost reacting with b-TCP. In the following 12 weeks little changeoccurred in the level of b-TCP relative to polymer, indicating theircomparable dissolution/degradation rates. After 12 weeks, how-ever, neither brushite nor monetite were readily detectable inthe samples, suggesting they had largely dissolved into the samplestorage water.

XRD studies were consistent with the Raman data but addition-ally enabled brushite and monetite to be more readily quantifiedand distinguished (see Fig. 2). With excess b-TCP, spectra weredominated by peaks (at 27.9, 31.2 and 34.5� 2h) due to this phaseas expected. With equimolar b-TCP and MCPM (peaks at 23 and24.3� 2h), the latter could readily be detected in dry samples butnot after 24 h in water. Instead, peaks due to both brushite (12�2h) and monetite (26.7 and 30.3� 2h) became visible. These de-clined relative to b-TCP peaks after 12 weeks in water, again con-sistent with faster dissolution of these more soluble calciumphosphates.

3.2. Water sorption and final mass loss

Water sorption of the polymer was about 4% (Table 1). That ofthe composites was between 8% when F% = 50% and T/M = 4 (i.e.,

lowest MCPM content) and 21% with F% = 70% and T/M = 1 (highestMCPM content). Factorial analysis (detailed in Section 2.9) showedthat water sorption was affected strongly by T/M (a2 = �0.6 ± 0.1),but less so by F% (a1 = 0.2 ± 0.02). The relative magnitudes andsigns of these ‘‘a” values are consistent with higher MCPM in thespecimens, encouraging greater water sorption. Final material losswas affected most by T/M (a2 = �1.0 ± 0.1), being 4% with T/M = 4and 10–12% with T/M = 1.

3.3. Compressive modulus

The average initial dry, 1-day wet and 70-day wet modulus ofthe polymer was 9, 7 and 3 MPa, respectively. All compositesexhibited higher modulus than the polymer at any given time(Fig. 3). The initial dry modulus for composites with F% = 70 and50% was 21–27 and 12–15 MPa, respectively. After 1 day in water,the modulus of all composite samples was 7–10 MPa and by70 days 5–8 MPa (Table 1).

0

5

10

15

20

25

30

0 20 40 60Time (days)

Com

pres

sive

mod

ulus

(MPa

)

0

10

20

30

0 1 2 3 4 5 6 7

Fig. 3. Changes in polymer and composite compressive modulus with time. s

polymer, j F% = 70%, T/M = 4, N F% = 70%, T/M = 1, X F% = 60%, T/M = 2, h F% = 50%, T/M = 4, 4 F% = 50%, T/M = 1. The results shown are averaged using Md of 90 and30 lm data. Error bars = ±SD, n = 6.

0

5

10

15

20

0 10 20 30 40SQRT of time (h0.5)

PO43-

rele

ase

(%)

b

a

3

4

5

6

7

8

9

Ca2+

rele

ase

(%)

850 X. Zhao et al. / Acta Biomaterialia 6 (2010) 845–855

Factorial analysis confirmed initial compressive modulus wasdependant primarily upon F% (a1 = 0.3 ± 0.04) with higher F% lead-ing to greater modulus. Additionally, reduction in modulus uponimmersion in water for 1 day was affected most by F%(a1 = 0.2 ± 0.02) but additionally by T/M (a2 = �0.1 ± 0.01). These‘‘a” values (see Section 2.9) indicated that an increase in F% and de-crease in T/M both resulted in greater reduction in modulus uponwater immersion.

0

1

2

0 10 20 30 40

SQRT of time (h0.5 )

Fig. 4. Cumulative phosphate (a) and calcium (b) release of composites as afunction of SQRT of time. s polymer, j F% = 70%, T/M = 4, N F% = 70%, T/M = 1, XF% = 60%, T/M = 2, h F% = 50%, T/M = 4, 4 F% = 50%, T/M = 1. The results shown areaveraged using Md of 90 and 30 lm data. Error bars = ±SD, n = 6.

3.4. Buffering capability of calcium phosphate fillers

3.4.1. Phosphate and calcium ion releaseBoth phosphate (Fig. 4a) and calcium (Fig. 4b) release profiles

exhibited burst release in the first 24 h (5 h0.5) followed by linearrelease vs. the square root (SQRT) of time for the remaining period.Cumulative release at 24 h and 12 weeks (45 h0.5) and the releaserate (slope of ‘‘release vs. SQRT of time” curve) between thesetimes are summarized in Table 2. Percentage phosphate releasewas typically between 2 and 3 times greater than percentage cal-cium release. Release also generally increased upon raising MCPMlevel. Cumulative phosphate release at 24 h and 12 weeks was 1–6and 8–22%, respectively. Calcium release was 0.6–2.6% and 4–9% atthese times.

Table 2PO3�

4 and Ca2+ release from composites and pH of the composite storage solution.

Formulations PO3�4

F% (%) T/M Md (lm) 24 h (%) 12 weeks (%) KPO4 (100%/h0.5)

70 4 90 1.8 ± 0.1 10.5 ± 0.1 27 ± 170 4 30 1.6 ± 0.1 10.3 ± 0.3 25 ± 170 1 90 5.9 ± 0.0 21.7 ± 0.7 35 ± 170 1 30 6.3 ± 0.2 19.4 ± 0.5 41 ± 360 2 60 2.5 ± 0.2 11.9 ± 0.3 23 ± 150 4 90 1.1 ± 0.2 8.2 ± 0.8 12 ± 150 4 30 1.1 ± 0.0 8.0 ± 0.5 12 ± 150 1 90 3.4 ± 0.2 19.0 ± 0.2 22 ± 150 1 30 2.5 ± 0.3 14.0 ± 0.0 23 ± 1

KPO4 and KCa are the gradient of the curve of ‘‘cumulative percentage of phosphate and(45 h0.5).* Average pH of polymer between 2–24 h and 24 h–12 weeks is 5.5 and 4.0, respectivel

From factorial analysis, the initial 24 h phosphate and calciumrelease percentage was affected on average most by T/M

Ca2+ Average pH*

24 h (%) 12 weeks (%) KCa (100%/h0.5) 2–24 h 24 h–12 weeks

0.9 ± 0.0 3.6 ± 0.8 8 ± 1 4.3 5.40.8 ± 0.1 3.9 ± 0.4 8 ± 1 4.2 5.52.6 ± 0.0 7.8 ± 0.6 13 ± 2 3.5 5.72.4 ± 0.2 7.3 ± 0.4 12 ± 4 3.5 5.71.2 ± 0.0 5.8 ± 0.1 10 ± 1 4.0 5.31.2 ± 0.1 3.7 ± 0.2 6 ± 1 4.2 4.60.6 ± 0.0 3.6 ± 0.8 7 ± 1 4.3 4.63.2 ± 0.1 8.5 ± 0.6 12 ± 1 3.7 5.41.8 ± 0.1 6.4 ± 0.1 9 ± 1 4.0 5.5

calcium release vs. SQRT of time”, respectively, between 24 h (5 h0.5) and 12 weeks

y. Data = average ± SD.

X. Zhao et al. / Acta Biomaterialia 6 (2010) 845–855 851

(a2(PO3�4 ) = �0.6 ± 0.1, a2(Ca2+) = �0.5 ± 0.1), then F% (a1(PO3�

4 ) =0.3 ± 0.02, a1(Ca2+) = 0.1 ± 0.01) and Md (a3(PO3�

4 ) = 0.1 ± 0.01,a3(Ca2+) = 0.2 ± 0.01) having much smaller effects. Decrease in T/M or increase in F%, leads to an increase in initial phosphate andcalcium ion release. Moreover, with larger Md more phosphateand calcium could be released. On average long-term release (seerelease constant K in Table 2), was also enhanced upon decreasein T/M or increase in F% (i.e., upon raising the level of MCPM inthe formulation).

3.4.2. pH measurementThe pH of the polymer storage solution decreased to on average

5.5 in the first 2–24 h and then fluctuated around 4.0 for theremaining 12 weeks. The average pH of the first 2–24 h storagesolutions for composites was lower with greater MCPM (i.e., lowT/M and high F%), being 3.5 with T/M = 1 and F% = 70% but between3.7 and 4.3 for all other composites. Conversely, the average pH inthe remaining 12-week period, was lower with less MCPM (i.e.,high T/M and low F%) being 4.6 with T/M = 4 and F% = 50% but be-tween 5.3 and 5.7 for all other composites (see Table 2).

3.5. Biocompatibility

3.5.1. Cell growth on polymer and composite surfacesThe number of cells present after 3 days of culture on the poly-

mer and composite films were all found to be >80% of that on thecontrol ThermanoxTM (Table 1). The cells cultured on all samplesexhibited an apparently normal morphology although the cells at-tached to the ThermanoxTM seemed to have the most pronouncedcytoplasmic processes (Fig. 5).

Fig. 5. SEM images of MG-63 cells incubated for 3 days on a polymer film (i), a compositM = 1 and Md = 30 lm (iii) and a ThermanoxTM plastic coverslip (iv). Note the granular suThermanoxTM (iv). The highly spread cells on the different material surfaces demonstrate

3.5.2. Tissue response to implanted materialsThe interaction between the bone and the implanted composite

material after 7 days of incubation is illustrated in Fig. 6. All fourfemurs implanted with composite films were found to be embed-ded in the CAM, with blood vessels surrounding the femur. Thecross-sectional diameter and the overall view of the femurs withthe implants (Fig. 6c) was comparable with the control sampleswithout any implant (Fig. 6a) and with the HA-implanted femurs(Fig. 6b). The implants appeared to be in close proximity to thebone, with no indication of an immunological reaction, e.g., nonoticeable accumulation of leucocytes, as shown by higher magni-fication in Fig. 6d. Bone adjacent to the implant site appeared his-tologically normal.

The CAM response to the control PLGA (Fig. 7a) was observed tobe comparable with the composite material (Fig. 7b), with bloodvessels appearing to be in close proximity to both the PLGA andthe composite. Moreover, blood vessel density proximal to boththe PLGA and the composite appeared similar to that in other areas,with no evidence of vessel hemorrhage in the region of the graftedmaterial.

4. Discussion

The adhesive materials investigated in this study can be in-jected into the injury site and light cured by a dental curing gunfor open wounds, or by an optical fibre for internal wounds. Alter-natively they may also be self-cured via chemical means using ben-zoyl peroxide as initiator and DMPT as accelerator [37].Additionally, in comparison with conventional PMMA bone ce-ments, these materials would release an order of magnitude less

e film with F% = 70%, T/M = 4 and Md = 30 lm (ii), a composite film with F% = 70%, T/rfaces of the composites (ii and iii) and the smooth surfaces of the polymer (i) andthe characteristic adhesive feature of active MG-63 cells. Scale bars: 100 lm.

Fig. 6. Histology of toluidine blue stained cross-sections of chick embryo femurs without (negative control; (a)) and with implantation with HA beads (positive control; (b))and composite films with F% = 50%, T/M = 1 and Md = 30 lm (c and d). Control and experimental femurs showed a similar bone structure. The outer surface of the femur wascovered by the CAM. The bone marrow (B), trabecular bone (T), periosteal tissue (PS) and the implant (HA and F) could all be readily identified. Note that the implant appearedto be in close proximity to the adjacent bone (arrows). Scale bars: 200 lm for (a–c); 100 lm for (d).

Fig. 7. CAM response to the PLGA microspheres (a) and implanted composite films (F) with F% = 50%, T/M = 1 and Md = 30 lm (b). Note the absence of negative reaction (e.g.,vessel hemorrhage) in the region of the grafted material (PLGA and F) and the presence of blood vessels (arrows) passing in close proximity to the implant. Scale bars: 1 mm.

852 X. Zhao et al. / Acta Biomaterialia 6 (2010) 845–855

heat upon curing because of their high monomer molecular weightand thus low methacrylate concentration. They are therefore unli-kely to cause thermal damage to any adjacent tissue.

In this study, chemical changes catalyzed by water sorption ofthe composites containing b-TCP/MCPM reactive fillers have beeninvestigated. The effects of F%, T/M and Md on the water sorption,elastic modulus and ion release kinetics have been quantified usinga factorial analysis experiment design. The in vitro and in vivo bio-compatibility of these materials has also been assessed.

Raman confirmed that the adhesive monomers could be fullypolymerized in the presence of calcium phosphate fillers. The levelof monomer conversion is greater than normally observed in acry-lates or methacrylates under comparable conditions [38], possiblybecause of the low glass transition temperature of the initialmonomers which would prevent the vitrification that limits finalpolymerization [15]. Such maximum monomer conversion ishighly desirable in degradable systems as it would limit the release

of reactive and potentially toxic double-bond-containing compo-nents upon degradation.

Raman with XRD additionally proved water sorption could cat-alyze conversion of b-TCP and MCPM to brushite and/or monetitewithin 24 h inside the polymer. The formation of both brushiteand monetite was consistent with previous studies [39]. Brushiteformation is kinetically favoured compared to monetite but itcan convert to monetite, particularly at low temperatures [40–42], in humid and slightly acidic conditions [43] (similar to reac-tion conditions in this study). The chemical composition of the cal-cium phosphates after 24 h water immersion was primarilymonetite and brushite or b-TCP and dependant on T/M. The ratioof produced brushite and monetite to b-TCP decreased in the orderT/M = 1 > T/M = 2 > T/M = 4 as expected.

Additionally, the relative amount of brushite and monetite topolymer decreased whereas the amount of b-TCP to polymer keptalmost constant during long-term water immersion. This was

X. Zhao et al. / Acta Biomaterialia 6 (2010) 845–855 853

probably due to the higher solubility of brushite and monetite thanb-TCP (solubility product constant (Ksp) at 25 �C of brushite, mone-tite and b-TCP = 2.59 � 10�7, 1 � 10�7 and 2.07 � 10�33, respec-tively [44]). This result was consistent with the higher phosphateand calcium release from specimens with T/M = 1 compared tospecimens with T/M = 4. The formation of brushite and monetite(T/M = 1) may be beneficial since the brushite or monetite fillersgenerated higher phosphate and calcium release associated withhigher pH compensation ability.

The addition of calcium phosphate particles enhanced the drycompressive modulus of the polymer by up to threefold. The de-gree of initial dry mechanical reinforcement depended mainly onthe amount of added filler. During the first 24 h of water immer-sion, composite modulus decreased far more than that of the poly-mer. This was probably due to the higher water sorption of thecomposites: the water ingress can act as a plasticizer, leading toswelling of material and reduction in material modulus [45]. Spec-imens with T/M = 1 were expected to exhibit largest decrease inmodulus due to highest water sorption. Surprisingly, the effect ofT/M on modulus reduction was less pronounced than F%. Thismay be due to production of brushite crystals binding some ofthe absorbed water which may reduce the possibility of polymerplasticization. Additionally the brushite/monetite may be morehomogeneously dispersed in the set polymer than the original par-ticles providing a novel mechanism of resisting modulus decline.

The modulus of the composites declined continuously but onlyvery slightly over the remaining 10 weeks, possibly because degra-dation of the polymer and dissolution of calcium phosphate wasvia surface erosion more than from the material bulk. In this perioddecline in modulus of the polymer was more pronounced, possiblydue to continuing bulk degradation. The dissolution of calciumphosphates within the polymer may also contribute to the mainte-nance of the modulus of the composite materials due to ‘‘salt-bridge” or ionic bond formation within the composite during thecourse of degradation. With increasing carboxylic acid groups (pro-duced during the polymer backbone degradation), the salt-bridgesstart to form between the carboxyl groups on polymer fragmentsand calcium cations from the calcium phosphates, resulting in‘‘an ionomer” whose ionic cross-links combined with partially de-graded polymer networks (still having relatively high molecularweight) may result in maintenance of compressive modulus [46–48].

The modulus of the composites obtained (10–30 MPa) was low-er than that of human trabecular bone (50–800 MPa) [49]. Suchelastic materials could, however, be used as temporary bone adhe-sives for infiltration/stabilization of, for example, osteoporoticbone that has become brittle due to the loss of collagen [50]. Asnew bone grows progressively against the receding surface of thedegrading material, and eventually integrates and replaces thematerial, the temporarily formed ‘‘bone-material” might be ex-pected to have increased modulus compared with the implantmaterial alone.

Composites with T/M = 1 generated higher phosphate and cal-cium release than those with T/M = 4, probably due to more solublebrushite and monetite formation in the specimens with T/M = 1.The ability to release phosphate and calcium ions may be beneficialin aiding bone mineralization and regeneration, as these ions haveshown to stimulate osteogenic differentiation [7]. In addition,through a dissolution–precipitation process, a bone-like minerallayer could develop, leading to new bone formation initiated bythe presence of osteogenic compounds in body fluids (e.g., bonemorphogenetic proteins). Moreover, the released ions could nucle-ate into the pores of bone collagen fibrils, resulting in increasedmineralization along the fibrils [51].

The PO3�4 /Ca2+ release ratio of 2–3 (ratio of KPO4 /KCa) for the

composites suggested that the phosphate and calcium elements

might be released in the form of Ca(H2PO4)2 with possibly somephosphoric acid (H3PO4). The release of acidic organic components,however, is likely to be very small as the total mass loss can be al-most entirely accounted for the release of calcium phosphates. Inthe presence of water, ion exchange with the produced monetiteor brushite could neutralize the acidic polymeric fragments form-ing monocalcium phosphate via the following reaction:

2RCOOHþ 2CaHPO4ðsÞ ! ðRCOO�Þ2Ca2þðsÞ þ Ca2þðH2PO�4 Þ2ðaqÞ

The initial pH drop with the composite specimens in the first24 h might be attributed to the release of higher level of phospho-ric acid [52]. The higher pH of the composite storage solution thanpolymer at later times indicates a buffering effect by the calciumphosphates particularly when T/M = 1, whereas when T/M = 4, thebuffering effect was less pronounced. This may be because thespecimens with T/M = 1 had a higher amount of more solublebrushite and monetite. The nature of the filler material is thus cru-cial for the long-term performance of an implant: their solubilityshould be suitable in order to provide sufficient pH compensation[53,54].

The results of phosphate and calcium ion levels as well as pH ofthe storage solution found above confirm that the fillers used pro-vide release of phosphate and calcium ions that could both aidbone regeneration and simultaneously, at least partly, buffer thereleased acidic polymer degradation products. Additionally, theoverall final amount of released acids and organic degradationproducts will be reduced by raising the filler level, as the fractionof polymer is lower.

The calcium phosphates present in the composite materialsmay be degraded via erosion, fragmentation and dissolution[43,55]. The results of this study suggest that the polymer de-graded slower than the calcium phosphates. Previous work hasshown that polymer discs of similar chemical structure to thoseused in this work could degrade completely in a few days in 1 MNaOH solution [8]. In Hepes buffered saline, however, degradationwas of the order of 1000 times slower. Both this earlier work andthe above new studies therefore suggest that the pure polymermay take several years to fully degrade in aqueous solution. Invivo, however, degradation may be accelerated by the presenceof enzymes [56] and cells [57]. The degradation products of theinorganic components are calcium and phosphate ions or particleswhereas the polymer degradation products are presumably polye-thers, monomeric and oligomeric lactic acid and oligo(methacylicacid). These are relatively nontoxic and can be excreted directlyor after entry and exit from various metabolic pathways [58].

The in vitro biocompatibility test indicated that bone-like MG-63 cells were found to attach and spread on both the polymerand composite surfaces, with an apparently normal morphologydespite the different surface nature of the polymer (smooth) andcomposites (granular). In these experiments, a 24 h period of pas-sivation was used to ensure the release and removal of any possi-bly deleterious products initially present in the materials. This wascarried out because, in other experiments, it was found that directcell seeding on the material surface without such passivation couldprevent cell adhesion and subsequent cell growth (data notshown). This may have been due to the faster initial material loss(e.g., release of small molecules), dissolution of the surface or unre-acted acidic MCPM (as shown by the result of initial pH decrease)or because of release of cytotoxic initiators [59]. However, the ini-tial release of possibly low levels of cytotoxic components may notpresent a serious problem in vivo, because of the rapid and contin-uous circulation of body fluid as well as the high buffering capacityof serum.

The formulation of F% = 50%, T/M = 1 and Md = 30 lm wasselected for the in vivo test due to its high phosphate and calcium

854 X. Zhao et al. / Acta Biomaterialia 6 (2010) 845–855

release and buffering capacity. Additionally, its fluidity facilitatesthe production of thin films. When composite films of this formu-lation were implanted into the chick embryo femur, they exhibitedclose contact with the adjacent bone, with no indication of adverseimmunological reaction. In addition, the response of the CAM tothe implant itself did not indicate any detectable vessel hemor-rhage in the region of the grafted material, and apparently normalblood vessels could be seen passing under the implant and in closeproximity to it. These observations indicate that the implantedcomposite material was unlikely to have been toxic to normal bonein vivo, and suggest that such materials could have potential forclinical application as bone and dental adhesives.

5. Conclusions

The injectable degradable composite materials investigated inthis study are elastic, release relatively high levels of phosphateand calcium ions, buffer acid production and appear to be biocom-patible. Increase in the amount of calcium phosphate fillers (irre-spective of b-TCP or MCPM) resulted in enhanced compressivemodulus of the dry materials although all composites exhibited acomparable modulus after water immersion. Composites with anequal molar ratio of b-TCP and MCPM (T/M = 1) formed relativelymore brushite and monetite after 24 h of water immersion, re-leased higher levels of phosphate and calcium ions and were thusmore capable of buffering acid production than those with excessb-TCP. All formulations tested were found to be biocompatible bythe procedure used here.

Acknowledgements

This research work was supported by the Engineering and Phys-ical Sciences Research Council, Medical Research Council and aDorothy Hodgkin Postgraduate Award. The authors thank Dr. NickyMordan for her valuable help with SEM, Dr. Sze Man Ho withmonomer synthesis and Drs. Tom Frenkiel and Geoff Kelly at theMedical Research Council Biomedical NMR Centre with NMRstudies.

Appendix A. Figures with essential colour discrimination

Certain figures in this article, particularly Figures 6 and 7, aredifficult to interpret in black and white. The full colour imagescan be found in the on-line version, at doi: 10.1016/j.actbio.2009.09.020).

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