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Biomaterials 29 (2008) 944–953

www.elsevier.com/locate/biomaterials

Osteoconduction and osteoinduction of low-temperature3D printed bioceramic implants

Pamela Habibovica,b,�,1, Uwe Gbureckc,1, Charles J. Doillond,e,David C. Bassetta, Clemens A. van Blitterswijkb, Jake E. Barraleta

aFaculty of Dentistry, McGill University, Strathcona Anatomy and Dentistry Building, 3640 University Street, Montreal, Quebec, Canada H3A 2B2bDepartment of Tissue Regeneration, Institute for Biomedical Technology, University of Twente, Zuidhorst Building,

Drienerlolaan 5, 7522 NB Enschede, The NetherlandscAbteilung fur Funktionswerkstoffe, ZMK-Klinik, Pleicherwall 2, 97070 Wurzburg, Germany

dDepartment of Surgery, Universite Laval, Quebec, Canada G1V 4G2eOncology and Molecular Endocrinology Research Centre, CHUL Research Centre, CHUQ, 2705 Boulevard Laurier,

Quebec City, Quebec, Canada G1V 4G2

Received 24 July 2007; accepted 17 October 2007

Available online 4 December 2007

Abstract

Rapid prototyping is a valuable implant production tool that enables the investigation of individual geometric parameters, such as

shape, porosity, pore size and permeability, on the biological performance of synthetic bone graft substitutes. In the present study,

we have employed low-temperature direct 3D printing to produce brushite and monetite implants with different geometries. Blocks

predominantly consisting of brushite with channels either open or closed to the exterior were implanted on the decorticated lumbar

transverse processes of goats for 12 weeks. In addition, similar blocks with closed channel geometry, consisting of either brushite or

monetite were implanted intramuscularly. The design of the channels allowed investigation of the effect of macropore geometry (open

and closed pores) and osteoinduction on bone formation orthotopically. Intramuscular implantation resulted in bone formation within

the channels of both monetite and brushite, indicating osteoinductivity of these resorbable materials. Inside the blocks mounted on the

transverse processes, initial channel shape did not seem to significantly influence the final amount of formed bone and osteoinduction was

suggested to contribute to bone formation.

r 2007 Elsevier Ltd. All rights reserved.

Keywords: Calcium phosphate cement; Three dimensional printing; Osteoconduction; Osteoinduction; In vivo test

1. Introduction

In the search for a synthetic biomaterial that is able tosuccessfully replace autografting, the ‘‘gold-standard’’ inorthopaedic and cranio-facial surgery, a range of materialshave been developed in the past four decades. Motivesinfluencing the design of synthetic bone graft substi-tutes are often based upon mimicking one or more

e front matter r 2007 Elsevier Ltd. All rights reserved.

omaterials.2007.10.023

ing author. Department of Tissue Regeneration, Institute

Technology, University of Twente, Zuidhorst Building,

, 7522 NB Enschede, The Netherlands.

9 3400; fax: +31 53 489 2150.

ess: p.habibovic@tnw.utwente.nl (P. Habibovic).

ibutions.

properties of natural bone, since this is the intended tissueto be repaired or augmented. Various types of calcium-phosphate biomaterials, which resemble either the compo-sition of bone mineral or its precursors have beendeveloped, such as hydroxyapatite-, a- and b-tricalciumphosphate, octacalcium phosphate and dicalcium phos-phate in the form of ceramics, cements and thin coatings[1]. Many relatively insoluble calcium-phosphate materialsare osteoconductive, and in some cases, even able to inducenew bone formation in extraskeletal sites, i.e. are osteoin-ductive [2–5].In the case of large or dense bone graft substitutes,

biodegradation is important for allowing simultaneousreplacement of the material with new bone. This process

ARTICLE IN PRESSP. Habibovic et al. / Biomaterials 29 (2008) 944–953 945

prevents stress shielding and disadvantageous resorption ofneighbouring bone. Although calcium-phosphate materialscan be degraded through cell-mediated processes (resorp-tion), chemical dissolution is the main pathway ofbiodegradation. Dissolution of calcium-phosphate materi-als is largely dependent on their chemical composition [6,7].For both dense and macroporous materials, the rate ofresorption should ideally be the same as the rate of newbone formation, in order to obtain a constantly stabilizedrepair and eventually a fully repaired bone defect withoutremnants of the synthetic material.

Apart from the physico-chemical properties of syntheticbone graft substitutes, geometry, ultrastructure, and there-fore mechanical properties are determinants for successfulbone healing of clinically relevant bone defects. Whenproduced in dense form, certain types of calcium-phos-phate ceramics have a sufficient compressive strength toprovide initial stability during the bone healing process,however, their fracture toughness is significantly inferior tothat of cortical bone and limits their use to non-loadbearing, or instrumented load bearing defects [8]. In orderto solve the problem of unpredictable dissolution of densebone graft substitutes and to mimic the structure ofcancellous bone, macroporosity is often introduced intosynthetic bone graft substitutes. Porosity can accelerate thebiodegradation of the material by increasing the surfacearea in contact with body fluids and enhancing thevascularization potential. Furthermore it may allow in-growth of bone tissue therefore improving the integrationof the implant. Various parameters of pore structure havebeen shown to be important for the performance ofsynthetic bone graft substitutes [9–12]. Classical methodsof introducing macroporosity into synthetic calcium-phosphate bone graft substitutes include foaming ofceramic slurries [13,14], dual phase mixing of polymerand ceramic slurries [15] and addition of organic particlesto ceramic powders [16,17]. However, control of para-meters such as porosity, pore size, pore interconnectivityand permeability is limited with these methods.

Free-form fabrication techniques for producing 3Dsynthetic bone graft substitutes, such as direct printing of3D implants from computer-aided design files, offers a toolfor precise control of the overall geometry and propertiesof the porous structure. In a recent paper, we havedemonstrated a direct process for 3D printing of calcium-phosphate structures at room temperature [18]. Wehave now employed this technique to produce brushite(dicalcium phosphate dehydrate, DCPD) cement implantswith precisely designed geometry in order to study theirbone forming capacity. The effect of implant shape onbone growth and material dissolution was investigatedusing an orthotopic decorticated lumbar transverse pro-cesses model in goats. Similar 3D printed implants,predominantly consisting of either brushite or monetite(dicalcium phosphate anhydrous, DCPA) were also im-planted ectopically in paraspinal muscles to study theirability to induce ectopic bone formation. In our previous

work, we have shown that increasing specific surface areaeither directly or indirectly by increasing the resorbabilityof calcium-phosphate ceramics, improves osteoinductivepotential [19]. Due to a lower water consumption, monetitestructures in general contain more micropores and thuspossess a higher specific surface area than brushitestructures produced using a similar method [20]; this wasthe rationale behind the choice of these two phases forintramuscular implantation.

2. Materials and methods

2.1. Study design

Twelve adult Dutch milk goats were used in this study after the

approval of the Utrecht University animal care committee (DEC, Utrecht

University, Utrecht, The Netherlands). For orthotopic implantation,

polymeric cassettes containing implants (Fig. 1a) were attached to the

right L4 transverse process. In addition to the materials described in this

paper, seven other cassettes containing implants were bilaterally attached

on the transverse processes of the L1–L4 lumbar vertebrae of each

goat, the results of which will be discussed separately. None of the

implanted materials contained cells or factors that could influence

neighbouring conditions. Surgical procedures were performed under

standard conditions.

To monitor bone formation over time, fluorochrome markers were

administered at 3, 6 and 9 weeks and the animals were sacrificed at 12

weeks after implantation. Bone formation was investigated by histology

and histomorphometry of non-decalcified sections using epifluorescent

and light microscopy.

2.2. Implant preparation and characterisation

The process of producing 3D printed brushite implants was described

previously [18]. In short, tricalcium phosphate (TCP) was synthesized by

heating a 2:1molar mixture of DCPA (CaHPO4, monetite) (Merck,

Darmstadt, Germany) and carbonated apatite (CC, CaCO3, calcite)

(Merck, Darmstadt, Germany) at 1400 1C for 14 h followed by quenching

to room temperature. The sintered cake was then crushed, sieved through

a 160mm sieve and milled in a planetary ball mill (PM 400, Retsch,

Germany) at 200 rpm with 500ml agate jars, four agate balls with a

diameter of 30mm, and a load of 125 g TCP per jar, for 30min.

Printing of cement implants was performed with a 3D-powder printing

system (Z-corporation, USA) using TCP powder and 20% phosphoric

acid as the liquid phase. Implant geometries were designed using CAD

software (Alibre DesignTM Xpress 9.2) and printing was performed

without anisotropic scaling (x ¼ y ¼ z ¼ 1.0). Printed implants were

stored in 20% phosphoric acid for 3� 60 s to increase the degree of

reaction towards DCPD. Monetite implants for intramuscular implanta-

tion were prepared by hydrothermal treatment at 134 1C (steam autoclave

sterilization) of printed brushite implants.

Brushite implants (blocks 6.5� 7� 8mm3) with two geometries, one

with nine open pore channels (+ 1.3mm, depth 8mm) (B 9OC) (Figs. 1b

and 1e) and one with nine closed pore channels (+ 1.3mm, depth 6mm)

(B 5+4CC) with openings on two opposite faces as shown in Figs. 1c and

1e were implanted in six goats each. For implantation in paraspinal

muscles, brushite and monetite implants (11.5� 8� 10mm3) with six

closed pore channels (+ 2.5mm, depth 8mm) (B 6CC and M 6CC) with

openings on one face (Figs. 1d and 1e) were used. Intramuscular

implantation of both B 6CC and M 6CC was performed in all 12 animals.

Prior to implantation, X-ray diffraction (XRD) patterns of the

materials were recorded using monochromatic CuKa radiation (D5005,

Siemens, Karlsruhe, Germany). Data were collected from 2y ¼ 201–401

with a step size of 0.021 and a normalized count time of 1 s/step. The phase

composition was checked by means of The International Centre for

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Fig. 1. Cage design with the locations of peripheral (X–X plane) and central (Y–Y plane) histological sections (a), design of B 9OC (b), B 5+4CC (c), B

6CC/M 6CC (d) with dimensions in mm and a digital photograph of the implants (e).

P. Habibovic et al. / Biomaterials 29 (2008) 944–953946

Diffraction Data reference patterns for a-TCP (PDF Ref. 09-0348),

b-TCP (PDF Ref. 09-0169), monetite (PDF Ref. 09-0080) and brushite

(PDF Ref. 09-0077). Post implantation, XRD patterns were recorded on

poly-methylmethacrylate (PMMA) embedded implants [21] using the

same method.

The porosity of the printed implants was calculated from the apparent

density after drying and the strut density calculated from the phase

composition and based on the density values from the literature [22]. The

gas adsorption analysis was used to determine the specific surface area of

the implants (TriStar 3000, Micromeritics, ASAP 200, USA). The

microstructure of different implants was characterized by using a scanning

electron microscope (SEM; JEOL 840A, Japan).

Mechanical testing was performed according to our own Standard

Operating Procedures which are based on ‘‘ISO 9917-1:2003: Dental

water-based cements. Powder/liquid acid–base cements’’. Mechanical test

samples were produced under the same conditions and had identical

microstructures to the implants; however, they did not contain channels.

Compressive strength testing was performed on cylindrical samples with a

height of 12mm and a diameter of 6mm under axial compression at a

crosshead speed of 1mm/min using a static mechanical testing device

(Zwick 1440, Ulm, Germany) and a 5 kN load cell. Samples for diametral

tensile strength (DTS) testing (h ¼ 5mm, d ¼ 10mm) were analyzed with

a cross-head speed of 10mm/min. The DTS was calculated from

DTS ðMPaÞ ¼ 2Fmax=ðpdhÞ, (1)

where d ¼ diameter (10mm), h ¼ height of the sample (5mm) and Fmax is

the failure load.

2.3. Implantation

All animals (about 2 years old, weighing 40–90kg), were housed at the

Central Animal Laboratory Institute (GDL), Utrecht, The Netherlands, at

least 4 weeks prior to surgery.

Cassettes used were as described previously [23,24]. The implants were

sterilized by Gamma irradiation at a dose of 20 kG (Isotron, Ede, The

Netherlands), while polyacetal components of the cassettes and the metal

screws were sterilized by autoclaving. Two implants (6.5� 7� 8mm3)

were placed into a cassette and separated by thin Teflon spacers

(0.5� 7� 8mm3).

Surgical procedures were performed under standard conditions [23,25].

After shaving and disinfecting the dorsal thoracolumbar area, a midline

ARTICLE IN PRESSP. Habibovic et al. / Biomaterials 29 (2008) 944–953 947

skin incision from T8-L5 was made to expose the paraspinal muscles that

were separated longitudinally to expose the transverse processes.

The process of the right L4 vertebrae was decorticated using an angled

bone rasp. One cassette, containing randomly located either B 9OC

or B 5+4CC implant, with five channel openings oriented toward

bone (Fig. 1a) and another implant (not discussed in this paper)

was screwed onto the process. Finger pressure was applied to the

top of the blocks in the cassette prior to muscle closure to ensure

direct contact of all blocks with the underlying bone as well as over-

lying muscle. After implantation on the transverse process, separate

intramuscular pockets were created in the T8-L3 paraspinal muscles and

filled with one B 6CC and one M 6CC implant. The fascia was then closed

with a non-resorbable suture to facilitate implant localization

at explantation. Durogesic 25 (fentanyl transdermal CII patches;

Janssen-Cilag EMEA, Beerse, Belgium) was administered for post-

operative pain relief.

Sequential fluorochrome markers were administered at 3 weeks

(CalceinGreen, 10mg/kg, Sigma, The Netherlands), 6 weeks (oxytetracy-

clin, engemycine 32mg/kg, Mycofarm, The Netherlands) and 9 weeks

(xylenol orange, 80mg/kg Sigma, The Netherlands) in order to visualize

bone growth dynamics.

At 12 weeks, the animals were euthanized by an overdose of

pentobarbital (Organon, Oss, The Netherlands). Spinal cassettes were

retrieved by cutting off the transverse processes, while intramuscular

implants were removed with some surrounding muscle tissue.

2.4. Histological processing and histomorphometry

After explantation, cassettes containing implants and underlying bone

were fixed in a solution of 5% glutaraldehyde and 4% paraformaldehyde

at 4 1C. Fixed samples were dehydrated by ethanol series (70–100%) and

transferred into a methylmethacrylate (MMA) solution that polymerized

at 37 1C within 1 week. Two longitudinal sections with a thickness of

10–15mm were cut through the middle of the three channel rows (one

central and two peripheral rows) as shown in Fig. 1a using a diamond saw

microtome (Leica Leitz 1600, Nussloch, Germany). One section was

stained with 1% methylene blue and 0.3% basic fuchsin after etching with

HCl/ethanol mixture for histology and one remained unstained for

epifluorescence microscopy of the fluorochrome markers.

All implanted materials were qualitatively analyzed using a light/

fluorescence microscope (Leica DMR, Germany). Histomorphometry was

performed on the transverse process implants only, because orientation of

the channels was impossible due to significant resorption of all

intramuscular implants. For histomorphometry, high-resolution digital

photographs (300 dpi) of the stained sections of B 9OC and B 5+4CC

implants were made. Prior to histomorphometrical analysis, bone and

material were pseudocoloured, green and red, respectively, using Adobe

Photoshops CS2. Image analysis was performed using a PC-based

system with KS400 software (version 3, Zeiss, Germany). The system

was geometrically calibrated with an image of a block of known

dimensions to analyze the area of the implant. A custom-made macro

was used to measure the area of interest, area of implant, area of bone,

implant outline available for bone apposition and contact length of bone

and implant. The measured histomorphometrical parameters were the

percentage of bone area in total implant area (%b), the percentage of

contact between bone and available implant outline in total implant area

(%b. cont) and the maximum distance between new bone and the

transverse process (%b height). In order to estimate the resorption of the

materials, the percentage of the scaffold area in the total implant area (%

scaff) was measured.

2.5. Statistics

Statistical calculations were performed using SPSS 14.0 software

(Chicago, IL). Due to large variations in the amount of bone between

individual animals, the data were not normally distributed in all cases.

Therefore, we chose the non-parametric Mann–Whitney test to compare

the two implant geometries after implantation on transverse processes,

using a significance level of po0.05.

3. Results

3.1. Materials characterization

Both brushite and monetite implants were faithfullyreproduced from hand-drawn CAD files (Fig. 1b–e).After post print hardening, the compressive strength ofthe brushite and monetite cements was 21.771.1 and8.370.4MPa, and the diametral tensile strength was7.470.7 and 1.470.2MPa, respectively. All implantscould easily be handled without fracturing.XRD patterns of brushite and monetite implants before

and after intramuscular and orthotopic implantation areshown in Fig. 2 displaying the presence of various calcium-phosphate phases. Before implantation, the brushitesamples were composed of a mixture of mainly brushitewith lesser amounts of monetite and unreacted TCP. Afterintramuscular implantation of the B 6CC the intensity ofthe brushite peaks strongly decreased while the monetiteand TCP peaks of both B 6CC and M 6CC were onlymarginally affected. In contrast, orthotopic implantationof the brushite samples resulted in the formation of aconsiderable amount of hydroxyapatite as a furthercrystalline phase from the hydrolysis of the secondaryphosphates monetite and brushite.Scanning electron microscopy was used to investigate the

microporous structure of brushite (Fig. 3a) and monetite(Fig. 3b) implants. The brushite surface consisted ofunreacted particles with a diameter of 10–20 mm, embeddedin a matrix of tabular crystals (2–5 mm). The monetitestructure was similar to that of brushite, however thesurface contained more micropores with a diameter of10–20 mm. The microporosity of brushite and monetite was38% and 44% and the specific surface area was 3.60 and6.09m2/g, respectively.

3.2. In vivo results

The original study design included 10 Dutch milk goats.5 weeks after surgery, one goat showed signs of inflamma-tory hoof disease (Laminitis) and was euthanized 1 weeklater after unsuccessful treatment. Immediately aftersurgery, another goat was treated with a single dose of anon-steroidal anti-inflammatory drug (NSAID, Diclofe-nacs) for similar Laminitis symptoms, after which thecondition improved. Although NSAIDs are well known fordisturbing the prostaglandin pathway and associated boneformation, we decided not to exclude this goat from theresults because the NSAIDs effect has only been describedfor longer periods [26]. Because of the illness of twoanimals, two new animals were included in the study. Thus,in total 12 animals were used for the study, and resultsfrom the 11 animals that survived the experiment arepresented.

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Fig. 2. X-ray diffraction patterns of powder printed brushite (left) and monetite (right) samples after 12-week intramuscular (IM) and orthotopic (OR)

implantation. The main peaks of brushite (B), monetite (M), X -tricalcium phosphate (T) and hydroxyapatite (H) are labelled.

Fig. 3. SEM micrographs of brushite (a) and monetite (b) surface structure prior to implantation. Scale bar ¼ 20 mm.

P. Habibovic et al. / Biomaterials 29 (2008) 944–953948

3.3. Orthotopic implantation

There were no signs of infection or other adversereactions associated with the implanted materials. Allcassettes were firmly attached to the underlying bone andsignificant ingrowth of new bone into the implant wasobserved in all cases. New bone was found in close contactwith the implant material, as well as inside the voids of thechannels in some animals (Fig. 4a). Considerable resorp-tion of both implants with open and with closed channels(B 9OC and B 5+4CC) was observed particularly in areaswith no or little bone formation (Fig. 4a). Investigation ofthe histological slides did not reveal the presence ofmultinucleated cells.

Analysis of fluorochromes showed that the 3-week labelwas only present in the lower third of the implants, closestto the host bone bed. The 6- and 9-week fluorochromelabels were found throughout the entire height of mostimplants regardless of channel geometry, demonstrating

bone apposition throughout the implant within 6 weeks(Fig. 4b). The pattern of fluorochrome labels indicated thatonce the entire internal channel surface was covered withbone (6 weeks), centripetal growth then followed (Fig. 4b).Histomorphometry was performed on histological sec-

tions cut across the width of the implants through theperipheral channel rows (X–X plane in Fig. 1a) and centralchannel rows (Y–Y plane in Fig. 1a) in order to determinethe effect of the initial area of channel openings in contactwith host bone bed on bone formation. Table 1 compareshistomorphometry between B 9OC and B5+4CC. While ofvalue for simple comparison, the data were found to be notnormally distributed. Therefore the boxplots (Fig. 5) moreaccurately represent the findings and the skewness of themeasurements. In peripheral sections, implants with openchannel design (B 9OC) showed higher values than theclosed channelled implants (B 5+4CC) for all measuredparameters, however, the difference between the twoimplant types was only significant for the percentage of

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Fig. 4. Digital micrographs of central sections of B 9OC (a, left) and B 5+4 CC (a, right), scale bar ¼ 1mm and of fluorescence microscope micrographs

taken at the top of B 9OC (b, left) and B 5+4 CC (b, right), scale bar ¼ 100mm, after 12-week implantation on lumbar transverse process. C ¼ implant

material; B ¼ new bone in (a) and level of bone after 12 weeks of implantation in (b); OTC ¼ 6-week fluorochrome label Oxytetracycline; XO ¼ 9-week

fluorochrome label Xylenol orange.

Table 1

Histomorphometrical data (mean7SD) of B 9OC and B 5+4CC measured on central and peripheral histological sections (Fig. 1a) and probability values

indicating statistical differences between B 9OC and B 5+4CC

Parameter B 9OC B 5+4CC p

% b central channel row 7.173.2 7.075.2 0.55

% b cont central channel row 6.975.2 8.275.8 1.00

mm b height central channel row 5.572.5 6.572.3 0.22

% scaff central channel row 13.2713.6 25.3710.7 0.15 (0.42 after correction)

% b peripheral channel rows 11.374.5 7.973.9 0.18

% b cont peripheral channel rows 12.474.3 6.673.4 0.012

mm b height peripheral channel rows 6.870.8 5.872.5 0.70

% scaff peripheral channel rows 28.4716.9 18.875.1 0.085 (0.035 after correction)

P. Habibovic et al. / Biomaterials 29 (2008) 944–953 949

bone contact (Table 1). In central sections, no significantdifferences between B 9OC and B 5+4CC were observedfor any of the histomorphometric parameters (Table 1).The percentage scaffold area relative to the total implantarea (% scaff) was comparable for the two implant types inboth central and peripheral implant sections. However,after correction for the initial difference in scaffold volume,B 5+4CC had significantly less material remnants than B9OC in the peripheral sections (Table 1).

3.4. Intramuscular implantation

There were no macroscopic or microscopic signs ofinfection and all implants were retrieved. Both brushite andmonetite implants had resorbed considerably; however, theamount of resorption varied from animal to animal.Ectopic bone formation was observed at different frequen-cies depending on implant material: in five out of 11 (45%)B 6CC, and in eight out of 11 (73%) M 6CC intramuscular

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%b %b.cont

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materialB 9OCB 5+4CC

materialB 9OCB 5+4CC

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Fig. 5. Boxplots showing histomorphometrical data of the percentage of bone occupying available pore area and percentage bone contact in the central

sections (left) and peripheral sections (right) of B 9OC and B 5+4CC after implantation on decorticated lumbar transverse processes. (�) outliers,

(1) extreme outliers.

P. Habibovic et al. / Biomaterials 29 (2008) 944–953950

implants. Qualitative histological assessment indicated thatmore bone was formed in monetite than in brushiteimplants (Fig. 6a). Bone formation was observed indifferent locations within the implant, but the largestamount of bone was found within the channels, close totheir openings (Fig. 6a). Generally, a layer of bone with athickness of about 100 mm was found in close contact withthe implant surface. Only in some cases had bone spreadinside the channel of the implant and covered a larger area.Contrary to the orthotopic implantation results, most boneformation was found in areas where the most materialresorption had occurred. The ectopically formed bone hadan immature, woven appearance with osteoblast layers,osteocytes and osteoid clearly visible (Fig. 6b). Fluores-cence microscopy revealed the presence of only the 9-weeklabel (Fig. 6c) for both brushite and monetite in someanimals, suggesting the start of bone formation betweenthe 6th and the 9th week of implantation. In other animalsthe onset of bone formation took place after the 9th week,based on the complete absence of fluorochrome labels.

4. Discussion

The use of free-form fabrication techniques, such asfused deposition modelling, 3D printing and 3D fibredeposition allows precise control of the parameters ofporous structures. Different rapid-prototyping techniqueshave been used for producing polymeric, sintered ceramicand metallic implants to study the influence of pore sizeand pore interconnectivity [24,27–30]. We have recentlypublished a method of producing calcium phosphateceramic structures by direct printing at room temperature,allowing for the production of thermally unstable phasesand for incorporation of biochemicals such as growthfactors and antibiotics [31]. In the present paper, osteo-conductive and osteoinductive properties of these 3Dprinted ceramics were investigated with an emphasis onthe effect of implant shape and pore geometry on new bone

formation. For this purpose, a decorticated lumbartransverse processes model in goats was used which allowssimultaneous comparison of multiple conditions [24,25].The two implant types were made of brushite. B 9OCimplants contained nine channels, open to both theunderlying bone of the transverse process and the overlyingparaspinal muscle, and was thus 100% permeable. Incontrast, B 5+4CC implants contained five channels opento the transverse process, but closed on the muscle side,and four channels with the reverse geometry. The totalvolume of the channels, i.e. the initial available space forbone growth, was about 25% lower in B 5+4CC than in B9OC implants. The most reliable histomorphometricparameter in this case would therefore be the percentageof bone in the available channel space. However, due to thehigh dissolution rate of the material and because it cannotbe assumed that the dissolution of both geometries wassimilar and linear with time, we decided to look at the areaof new bone formed in the region of interest, i.e. boneformation in the total implant area. Both implantgeometries showed filling of the channels with new bonethat initiated at the transverse process and sometimesreached the submuscular end of the channels. Fluorescencelabels showed that at 6 weeks, a layer of new bone hadalready formed on the walls of the channels over the totalchannel length, after which the bone formation continuedtowards the centre of the channel voids. As we did notobserve significant differences in the amount of bone or inthe fluorochrome label patterns between the open andclosed channel geometry, open macroporosity did not seemto be essential for bone ingrowth. In an earlier study, inwhich solid brushite implants produced in a similar waywere implanted into the femoral extensor muscles ofSprague-Dawley rats, a slower dissolution of the materialwas observed [32]. From these results it is reasonableto conclude that the resorption behaviour and thephase change of the material are dependent on the modelused both in terms of species and implantation site.

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Fig. 6. Digital micrographs of histological slides of B 6CC (a, left) and M 6CC (a, right) scale bar ¼ 1mm, light microscope (magnification 20� )

micrographs of B 6CC (b, left) and M 6CC (b, right), scale bar ¼ 50 mm and fluorescence microscope micrograph of M 6CC (c), scale bar ¼ 100mm after

12-week intramuscular implantation. C ¼ implant material; B ¼ new bone in (a) and (b) and level of bone after 12 weeks of implantation in (c); XO ¼

9-week fluorochrome label Xylenol orange.

P. Habibovic et al. / Biomaterials 29 (2008) 944–953 951

Indeed, in our study, orthotopically, but not intramuscu-larly, implantation of brushite led to a partial transforma-tion into HA. This is probably due to lower fluid exchangein orthotopic defects compared to soft tissue. This is a well-known problem preventing continuous degradation ofbrushite cements as the DCPD component in brushitecement may, depending on conditions, be stable, dissolveor hydrolyze to HA [33]. Short-term implantation studiesare required in order to gain a better insight into theresorbative properties of these materials and the effect onbone formation and possible bone remodelling.

With the design of B 5+4CC implants we also aimed atinvestigating the effect of tissue type on bone formation.A few groups have shown that certain calcium phosphatesare able to induce bone formation ectopically as recentlyreviewed [3]. It has furthermore been demonstrated that

such osteoinductive implants perform better in clinicallyrelevant critical size orthotopic defects [34]. By designingthe implants in such a way that certain channels were opento bone on one side and closed to muscle on the other, onlyosteoconduction of the material is investigated, while withthe inverse channel design, the effect of underlying bone isexcluded and only the ability of the material to induce newbone formation when in contact with overlaying muscle isinvestigated. In other words, both osteoconductive andosteoinductive properties can be investigated simulta-neously. It should however be noted that although thechannels of the B 5+4CC implants were closed on oneside, the material itself was microporous, so the implantwas not completely impermeable. Due to the minutemicropore size (maximum 20 mm) it is highly unlikely thatthere was a direct cell effect of the decorticated transverse

ARTICLE IN PRESSP. Habibovic et al. / Biomaterials 29 (2008) 944–953952

process on the bone formation in channels open to muscletissue. Indeed, we cannot exclude the effect of endogenouscytokines present in bone tissue or secreted by osteogeniccells of the transverse process on bone formation on theside of the muscle tissue. Further optimization of themodel, by using a 100% dense material for example, or bynot decorticating transverse process would give conclusiveevidence for the effect of one tissue on the other. We didnot find a significant difference in the amount of boneformed between the central and the peripheral sections(Fig. 1a), indicating that smaller initial contact with bonein this case did not have an effect on the final amount ofbone formed which may suggest an effect of osteoinduction,i.e. initiation of bone formation from the overlayingmuscle. However a solid conclusion is impossible giventhe unexpectedly high dissolution rate of the materials. Toinvestigate the separate effects of bone and muscle on newbone formation, earlier explantation time points or the useof a less degradable material, such as HA would be useful.

Implantation in paraspinal muscles nevertheless demon-strated for the first time that both brushite and monetiteprinted implants were able to induce ectopic boneformation, i.e. were osteoinductive. Gosain and colleagues[35] implanted apatite-based cements in sheep soft tissueand observed bone formation after 1 year of implantation.In a study by Yuan and coworkers [36], osteoinduction wasobserved after intramuscular implantation of prehardenedapatite cements in dogs, however, no information wasgiven on the degradation of the implanted materials andrelatively few samples were examined. This is the first timethat osteoinduction in goats is reported for degradablebrushite and monetite bioceramics. In a recent study byBodde et.al. [37], preset porous calcium phosphate cementconsisting of a mixture of TCP, monetite, calciumcarbonate and HA was implanted subcutaneously in goatsfor 3 and 6 months, and no bone formation was found ateither time point. The authors suggested that a fastdegradation of the cement was the main reason for thelack of ectopic bone formation. In our study, boneformation occurred on the surface of the cement in thecrevices formed due to dissolution of the material.Interestingly, although not exclusively, most bone forma-tion was found in the interior of the peripheral channels, ata relatively short distance (1–2mm) from the channelopenings. Observation of the histological sections indicatedthat this was also the location with the most pronouncedresorption of the material. As suggested in our previouswork, this indicates that a relatively ‘‘reactive’’ surface ofthe material is needed for initiating ectopic bone formationand that bone preferentially forms inside the ‘‘protected’’areas such as pores and channels, rather than on theimplant periphery [38]. The fact that bone had formed closeto the channel openings rather than deep inside thechannels might either be a direct effect of a better nutrientand oxygen supply or an indirect effect of the morepronounced dissolution/reprecipitation activity of theimplant surface. In our earlier work we have shown that

although surface reactivity positively influences ectopicbone formation, a relatively stable surface is needed tofacilitate new bone growth [38]. The dissolution rate of thematerials in this study was high, particularly in the areaswhere most bone formation was observed. Based on thefluorochrome labels, the time at which bone startedforming was 6–9 weeks in some and after 9 weeks in otheranimals. Nine weeks is rather late, as in similar modelsbone was induced between the 3rd and the 6th week bybiphasic calcium-phosphate ceramics [34,38,39]. A possibleexplanation for the observed bone formation despite therapid dissolution of the material could lie in the timing ofbone formation. It is conceivable that initially thedissolution of the surface was too high due to rapid fluidexchange preventing hydrolysis to HA or bone formationto occur. After a longer period of time, supersaturation ofthe local environment with calcium and phosphate ionsoccurred, stabilizing the surface with new calcium-phos-phate layer and facilitating bone formation.

5. Conclusion

In the present in vivo study we have used low-tempera-ture direct 3D printing to produce brushite and monetiteimplants with controlled geometry. In a decorticatedtransverse process model, these implants were used toinvestigate the effect of open and closed macroporosity onnew bone formation. Bespoke implants were designed insuch a way to create both ectopic and orthotopic cellularenvironments at the implantation site of transverseprocesses, allowing for the simultaneous investigation intothe relevance of osteoinduction and osteoconduction inbone repair. Initial open porosity of the implants did notseem to significantly influence the final amount of newbone formation. Intramuscular implantation of bothbrushite and monetite ceramics demonstrated their os-teoinductive potential and the effect of osteoinductivitywas also suggested orthotopically.

Acknowledgements

Authors thank Remi Tibben for technical assistance andthe Department of Orthopedics of the University MedicalCenter Utrecht, The Netherlands, for surgery and criticalreading of the manuscript. This study was financiallysupported by Canada Research Chair program (J.E.B.),and Quebec Ministere des Relations Internationales(Quebec-Bavaria Exchange Program) PSR-SIIRI-029 IBI.

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