Bioactive and Mechanically Strong Bioglass�-Poly(D,L-Lactic Acid)Composite Coatings on Surgical Sutures

Q. Z. Chen, J. J. Blaker, A. R. Boccaccini

Department of Materials and Centre for Tissue Engineering and Regenerative Medicine, Imperial College London,London SW7 2BP, United Kingdom

Received 5 March 2005; revised 30 April 2005; accepted 10 May 2005Published online 13 September 2005 in Wiley InterScience (www.interscience.wiley.com). DOI: 10.1002/jbm.b.30379

Abstract: New coating processes have been investigated for degradable (Vicryl�) and non-degradable (Mersilk�) sutures with the aim to develop Bioglass� coated polymer fibers forwound healing and tissue engineering scaffold applications. First, the aqueous phase of aBioglass� particle slurry was replaced with a poly(D,L-lactic acid) (PDLLA) polymer dissolvedin solvent dimethyle carbonate (DMC) to act as third phase. SEM observations indicated thatthis alteration significantly improved the homogeneity of the coatings. Second, a new coatingstrategy involving two steps was developed: the sutures were first coated with a Bioglass�–PDLLA composite film followed by a second PDLLA coating. This two-step process of coatinghas addressed the problem of poor adherence of Bioglass� particles on suture surfaces. Thecoated sutures were knotted to determine qualitatively the mechanical integrity of the coat-ings. The results indicated that adhesion strength of coatings obtained by the two-step methodwas remarkably enhanced. A comparative assessment of the bioactivity of one-step andtwo-step produced coatings was carried out in vitro using acellular simulated body fluid (SBF)for up to 28 days. Coatings produced by the two-step process were found to have similarbioactivity as the one-step produced coatings. The novel Bioglass�/PDLLA/Vicryl� and Bio-glass�/PDLLA/Mersilk� composite sutures are promising bioactive materials for woundhealing and tissue engineering applications. © 2005 Wiley Periodicals, Inc. J Biomed Mater Res PartB: Appl Biomater 76B: 354–363, 2006

Keywords: surgical sutures; Bioglass�; poly(D,L-lactic acid); coatings; bioactivity; mechan-ical adherence

INTRODUCTION

The combination of bioresorbable polymers and bioactiveinorganic materials, such as calcium phosphates and bioac-tive glasses, to form biocomposites is being increasinglyexplored for biomedical applications, including wound heal-ing, bone reconstruction, and tissue engineering.1–4 One ofthese attempts has been to coat surgical meshes and sutureswith bioactive glasses.5–7 Synthetic bioabsorbable suturesbased on polyglycolic acid (PGA), polylactic acid (PLA), andpolyglycolide lactide (PGLA) copolymers have been studiedextensively for more than 30 years.8–14 In particular thecomposite polyglactin 910 has shown excellent performanceas sutures in wound healing applications.8,9 These suturesexhibit three important advantages: minimal tissue reactions,good mechanical properties, and easy and reproducible fab-rication.8 Thus this fibrous material is considered a promisingprecursor to fabricate resorbable three-dimensional (3D) scaf-

folds incorporating bioactive ceramics or glasses for tissueengineering applications.5–7

The rationale of coating a polymeric scaffold with bioac-tive ceramics is that the bonding strength of the bioresorbablepolymer to bone and other tissue can be enhanced by thepresence of the bioactive ceramic coatings.5 It has beenshown that bioactive glasses can form strong bonds to boneand soft tissues.15,16 This bonding capability to living tissuesis referred to as bioactivity and has been associated with theformation of a carbonated hydroxyapatite layer on the glasssurface when implanted or in contact with biological fluids.15

In previous work, Bioglass� 45S5 particles were coated onpolyglactin 910 sutures using aqueous slurry dipping as theprocessing method.5,6 The coated sutures exhibited a highlevel of chemical reactivity, that is, carbonated hydroxyapa-tite formation in simulated body fluid (SBF), indicating theirbonding potential to bone tissue. However, the bioactivitywas hampered by the poor adherence of the 45S5 Bioglass�coatings to the sutures. Obviously, the coatings would con-tribute nothing to the bonding of tissue to scaffolds if theypeeled off from the scaffolds, no matter how strongly theycould bond to bone.17 An approach to tackle this issue, as

Correspondence to: A. R. Boccaccini (e-mail: a.boccaccini@imperial.ac.uk)

© 2005 Wiley Periodicals, Inc.

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followed in this study, is to introduce another component inthe coating slurry, which will act as glue promoting theadhesion of the glass particles to the suture and mesh sur-faces. Poly(D,L-lactic acid) (PDLLA) was chosen for thisfunction. This biodegradable polymer has been extensivelyinvestigated as a biomedical coating material because of itsexcellent features with respect to implant coating perfor-mance.18,19 It exhibits high mechanical adherence to sub-strates,20 good osteoinductive potential, and excellent bio-compatibility in vivo.21 The material also shows good anti-thrombogenic characteristics.22

In the present work, a new method is developed to coatcommercially available surgical sutures with Bioglass�–PDLLA composite films using a combined technique com-posed of slurry-dipping and slurry-spraying. The microstruc-ture of the coatings was investigated by scanning electronmicroscopy (SEM), while the mechanical stability of thecoatings was qualitatively determined by knotting the suturesand observing the resultant coating morphology and micro-structure. The effect of the PDLLA/Bioglasss� coating on thebioactivity of the sutures was assessed by immersion in SBF.

MATERIALS AND EXPERIMENTAL PROCEDURES

Sutures

The sutures used in this work were violet braided biodegrad-able 3/0 coated Vicryl� (Polyglactin 910) and black non-resorbable 2/0 Mesilk� sutures produced by Ethicon Inc.(Edinburgh, Scotland). Vicryl� is a clinically used suturemade from the semi-crystalline poly(glycolide-L-lactide)(PGLA) copolymer of 90:10 mole ratio.8 The mean diameterof as-received Vicryl� suture is 0.33 mm. The coated Vicryl�material is treated with Polyglactin 370 and calcium stearatefor lubrication to improve its passage through tissue, knotplacement, and tie down. The Mersilk� suture used here is abraided silk suture with a mean diameter of 0.43 mm, and iscoated with beeswax.7

Preparation of Coating Slurries

The starting materials for fabrication of bioactive coatingswere PDLLA (Purasorb�, Purac Biochem, Gorinchem, TheNetherlands) and melt-derived 45S5 Bioglass� powder (par-ticle size � 5 �m).23 Purasorb� PDLLA with an inherentviscosity of 1.62 dL/g was completely amorphous and used

without further purification. Dimethyle carbonate (DMC)of � 99% purity was used as solvent, following positiveresults of our previous investigations.24 The preparation andcomposition of slurries for coatings are summarized inTable I.

Stable slurries containing three different amounts of Bio-glass� particles (20, 30, 40 wt %) were prepared as follows:PDLLA was dissolved in DMC at the ratio of 5 g PDLLA per100 mL DMC, by stirring for 2 h; then given amounts ofBioglass� powder were added to the solution of PDLLA andDMC, followed by vigorous stirring for 2 h.

Coating Methods and Processes

Table II lists the two coating processes used in this work. Inthe one-step coating process, a Bioglass�–PDLLA compositelayer was coated on sutures using slurry-dipping technique.The sutures were first immersed in DMC for 2 days in orderto assess whether DMC has any effects on suture integrity.SEM observations indicated that no considerable morpholog-ical changes occurred to the sutures after expose to DMC for2 days. The dipping time used in the one-step coating methodwas 5 min. In the two-step process, the coating procedurecarried out in the one-step process was followed by applica-tion of a second PDLLA coating. The second PDLLA coatingwas applied by spraying (using a commercial perform bottle)rather than dipping because the composite coatings of the firststep would be redissolved quickly into the solution if thedipping method has been used. The spraying method wasbased on a PDLLA–DMC coating solution of the followingcomposition: 5 g PDLLA per 100 mL DMC. The sutureswere rotated manually while being sprayed. The sprayingtime was about 10 s. After the coating procedure, the sutureswere dried in air at room temperature for at least 12 h.

SBF Assessment

This part of the study was carried out with the use of thestandard acellular in vitro procedure described by Kokubo

TABLE I. Preparation of Coating Slurries

Coating Preparation

Dissolved PDLLA PDLLA dissolved in DMC at a ratio: 5g PDLLA per 100 mL DMCPDLLA–Bioglass� slurry 20 wt % Bioglass�

Dissolved PDLLA � 30 wt % Bioglass�40 wt % Bioglass�

�wt % of the total weight of the slurry*

* Total weight of the slurry � weight of PDLLA � weight of DMC � weight of Bioglass�.

TABLE II. Coating Procedures

One-Step CoatingBioglass�–PDLLA composite coating by slurry-dipping.

Two-Step CoatingBioglass�–PDLLA composite coating by slurry-dipping,followed by PDLLA coating applied by spraying after thecoatings obtained in the first step were dried.

355COMPOSITE COATINGS ON SURGICAL SUTURES

and colleagues.25 The sutures were immersed in 75 mL ofSBF in clean conical flasks, which had previously beenwashed using HCl and deionized water. The conical flaskswere placed inside an orbital shaker (New Brunswick Scien-tific, Classic Series, C24 Incubator Shaker), which rotates at175 rpm at controlled temperature of 37°C. The pH of thesolution was maintained constant at 7.25. Three samples of 3cm length were extracted from the SBF solution after giventimes of 4, 7, 14, and 28 days. The SBF was replaced twicea week because the cation concentration decreased during thecourse of the experiments, as a result of the changes in thechemistry of the samples, as explained below. Once removedfrom the incubation flasks, the samples were rinsed gentlyfirst in pure ethanol and then with deionized water and left todry at ambient temperature.

Characterization

The microstructure of selected coated sutures was character-ized in a JEOL 5610LV scanning electron microscope (SEM)

before and after SBF immersion. Samples were gold-coatedand observed at an accelerating voltage of 15 to 20 kV. Somesutures were also characterised with the use of X-ray diffrac-tion (XRD) analysis with the aim to assess the formation ofhydroxyapatite crystals on material surfaces after differenttimes of immersion in SBF. A Philips PW 1700 Seriesautomated powder diffractometer was used, employingCuK� radiation (at 40 kV and 40 mA) with a secondarycrystal monochromator. Data were collected over the range of2� � 5° to 100° using a step size of 0.04 and a counting timeof 25 s per step. The adhesion of coatings to sutures wasassessed qualitatively by making knots and investigating themacroscopic appearance and microstructure of the coatingsunder SEM.

RESULTS AND DISCUSSION

No significant difference was observed between the coatingson Vicryl� and Mersilk� sutures. Hence, in this article results

Figure 1. SEM micrographs of Bioglass�/PDLLA coated Vicryl� sutures. Coatings were producedafter immersion for 1 min in the slurry of (a, b) 20 wt % Bioglass�–PDLLA and (c, d) 40 wt %Bioglass�–PDLLA, using DMC as solvent.

356 CHEN ET AL.

will be shown and discussed using either Vicryl� or Mersilk�sutures, and the term sutures has been used throughout thissection to refer to the two suture types.

Coatings Produced by the One-Step Process

Initial qualitative analysis of the morphology and uniformityof the PDLLA–Bioglass� coatings was conducted by simple

visual inspection. The analysis of the coatings microstructurewas carried out by SEM observations. It was found that theimmersion time did not have considerable influence on themicroscopic characteristics of the coatings. Instead it was theconcentration of Bioglass� particles in the slurries that influ-enced the coating thickness and uniformity, as shown inFigure 1. With increasing Bioglass� content in the slurry, thehomogeneity of the coatings was significantly improved. Thecoatings exhibiting highest homogeneity were achieved witha slurry of 40 wt % Bioglass� for the Vicryl� sutures, asobserved on SEM micrographs in Figure 1(c,d); and 30 wt %for the Mersilk� sutures. It was, however, apparent that thespeed of removal of sutures from the slurry affected thehomogeneity of the coatings. Homogeneous coatings wereobtained at a withdrawal velocity of �5 cm/s. Droplets of thecomposite slurry could form on the sutures if they wereextracted too slowly.

The coated sutures using 30 wt % and 40 wt % Bioglass�–PDLLA slurries were knotted to qualitatively determine themechanical integrity of the coatings and their adhesionstrength to the suture substrate. Preliminary results indicatedthat adhesion of the composite coatings was improved, com-pared with pure Bioglass� coatings obtained by slurry dip-ping on similar sutures in our previous investigations.5–7 Thisimprovement can be attributed to the addition of the PDLLAmatrix in the coating, which should bond strongly to thesuture surface. However, indication of peeling and detach-ment of the coating layer was still evident on both Vicryl�and Mersilk� coated sutures, in particular at the two ends ofthe knots where the sutures suffered severe deformation andfriction, as shown in Figure 2.

Figure 2. SEM micrograph of a 40 wt % Bioglass�–PDLLA coated Vicryl� suture that has beenknotted. The coating peeled off at the two ends of the knot where the suture suffered severedeformation and friction.

Figure 3. Schematic graphs of coating morphologies. (1) Coating ofpure glass particles, as produced in previous studies5–7; (b) glass–polymer composite coatings achieved by one-step procedure; and (c)glass–polymer composite coatings achieved by two-step coatingprocedure.

357COMPOSITE COATINGS ON SURGICAL SUTURES

The resultant typical morphologies of pure Bioglass� andBioglass�–PDLLA coatings are schematically illustrated inFigure 3(a,b), respectively. It was expected however thatPDLLA would function as glue for Bioglass� particles, act-ing as a continuous matrix as indicated in Figure 3(c). Figure1(b,d) indicate that, on the contrary, in coatings fabricated bythe one-step method, the Bioglass� particles were not fullyembedded in the PDLLA matrix. The particles were in factadhered on the surface of a thin polymer layer and onlypartially embedded in it, as schematically shown in Figure3(b). The characteristics shown in Figure 3(b) actually indi-cate the possible coating mechanism of these Bioglass�–PDLLA composite films; the dissolved PDLLA developspolymer/polymer bonds in contact with the suture surface ata faster rate than the bonding of the suture to the inorganicglass phase. Thus, Bioglass� particles remain mainly attachedonly to the thin PDLLA film.

Coatings Prepared by the Two-Step Process

To achieve the coating structure shown in Figure 3(c), asecond PDLLA coating was applied. The SEM morphologiesof these two-step produced coatings (Figure 4) indicate thatthe desired characteristics shown in Figure 3(c) wereachieved, that is, Bioglass� particles were well embedded inthe second PDLLA layer. The mean diameter of the coatedVicryl� sutures was measured to be 0.57 � 0.03 mm, whichindicates that the coating thickness is �0.12 mm. The uni-formity of the coating was assessed qualitatively by SEMobservation of coated sutures and this was confirmed alongthe length (3 cm) of coated sutures.

These two-step coated sutures were knotted and examinedunder the SEM. Figure 5 shows that the sutures retained theircoating structure after being knotted. There were almost nocracks in the coating of the knotted suture, as shown in Figure5. The results were similar to both Vicryl� and Mersilk�sutures. They indicate that a first coating using a 30 to 40 wt% Bioglass�–PDLLA film followed by a second coating ofPDLLA thin film represents an optimum protocol for bioac-tive coating of sutures in terms of microstructural homoge-neity, adhesion strength, and mechanical integrity.

The mechanical properties of the coated sutures were notdetermined in this work. The measurement of the tensilestrength of sutures coated by a dry-powder method and by theslurry-dipping method has been presented elsewhere.6,7,26 Itwas found that Bioglass� coating on Vicryl� sutures by a drypowder pressing technique led to a slight reduction of su-tures’ tensile strength (463 to 404 MPa), which was ascribedto mechanical damage introduced on the suture surfacesduring the pressing of the hard glass particles. The coatingtechnique developed in the present study based on incorpo-rating a third phase (PDLLA) represents a less aggressiveapplication of the Bioglass� coating due to the presence ofthe viscous PDLLA and it is thus unlikely to induce mechan-ical damage during processing.

Figure 4. SEM micrographs of coated Mersilk� sutures. (a) 30 wt % Bioglass�–PDLLA compositecoating followed by PDLLA coating; (b) 40 wt % Bioglass�–PDLLA composite coating followed byPDLLA coating.

Figure 5. SEM micrograph of a two-step coated Vicryl� suture thathas been knotted, showing high structural stability of the PDLLA/Bioglass� coating layer.

358 CHEN ET AL.

Figure 6. Surface morphologies of one-step coated Vicryl� sutures after immersion in simulated bodyfluid for (a) 4 days, (b) 7 days, (c) 14 days, and (d) 28 days. The coating slurry was 40 wt % Bioglass�

in PDLLA–DMC solution.

Figure 7. XRD patterns of one-step coated Vicryl� sutures after immersion in SBF for 4, 7, 14, and 28days, showing HA formation. The spectrum of an as-received Vicryl� suture is given for comparison.The coating slurry was 40 wt % Bioglass� in PDLLA–DMC solution. [Color figure can be viewed in theonline issue, which is available at www.interscience.wiley.com.]

359COMPOSITE COATINGS ON SURGICAL SUTURES

Figure 8. Microstructures of the bonelike crystalline apatite on coated Vicryl� sutures after immersionin simulated body fluid for 2 weeks. (a) Fine apatite particles formed on a one-step produced coatingand (b) very fine, needlelike, apatite crystals on a two-step produced coating. The coating slurry of thefirst step was 40 wt % Bioglass� in PDLLA–DMC solution.

Figure 9. Surface morphologies of two-step coated Vicryl� sutures after immersion in simulated bodyfluid for (a) 4 days, (b) 7 days, (c) 14 days, and (d) 28 days, showing HA formation. The coating slurryof the first step was 40 wt % Bioglass� in PDLLA–DMC solution.

360 CHEN ET AL.

In Vitro Bioactivity of Coated Sutures in SBF

The surface reactivity of coated sutures was examined inacellular SBF as a qualitative indication of their in vitrobioactivity, following procedures of current practice in thebiomaterials field.25 The goal was to study the formation ofhydroxyapatite (bonelike apatite) on the surface of coatedsutures during immersion in acellular SBF, which is related tothe bioactivity of the materials.15,23,25 In particular, it is ofsignificance to carry out a comparative assessment of one-step and two-step produced coatings in order to investigatewhether or not the PDLLA film deposited in the two-stepcoating process has any negative effect on the bioactivity ofthe Bioglass�/PDLLA coatings.

Figure 6 shows the surface morphologies of one-stepcoated Vicryl� sutures after immersion in SBF for 4, 7, 14,and 28 days. XRD analysis (Figure 7) indicated that hydroxy-apatite was already formed on the coatings after immersionfor 4 days, and its amount apparently increased with time, asdocumented in Figures 6 and 7. A closer observation revealedthat the apatite layer contained fine particles that were � 0.2�m in size, as illustrated in Figure 8(a). For comparison,Figure 8(b) shows the microstructure of a two-step depositedcoating on a Vicryl� suture after immersion in SBF for 2weeks.

The reaction of the two-step coatings upon immersion inSBF began with the formation of porosity on the PDLLAlayer, as shown in Figure 9(a). Although apatite was notobserved on the surface by SEM at this stage (i.e., 4-day SBFimmersion), XRD already clearly indicated the existence ofhydroxyapatite, as shown in Figure 10. In addition, the in-tensity of HA peaks was as strong in the two-step coating asin the one-step coating (compare Figure 10 with Figure 7). Itis apparent that SBF could flow through the pores in the outerPDLLA layer to reach Bioglass� particles underneath the

surface and thus the formation of hydroxyapatite proceededin a similar way to that in Bioglass� coated samples.5–7 Theperforated thin PDLLA layer, therefore, plays an importantrole in the optimized behavior of the present coated sutures;it maintains the mechanical integrity of the coating, whileproviding microchannels for the communication betweenSBF and Bioglass� particles.

The increasing amount of HA crystals (bonelike crystal-line apatite) was clearly detected by SEM observations afterimmersion for 7 days and longer, as seen in Figure 9(b,d).Hydroxyapatite was recognized by its globular, cauliflowershape. A high magnification observation showed that eachglobular structure was made up of very fine apatite needles,which were about 0.25 �m in length and 0.05 �m in diam-eter, as illustrated in Figure 8(b). This shape of HA formed onBioglass� surfaces has been observed by several research-ers.15,16,27,28

Figure 9(c,d) shows also that individual globular forma-tions grew and linked together in a later stage (� 14 daysimmersion) to form an apatite layer covering the surface ofthe coated suture. However, the formation of bonelike apatitewas not homogeneous on the surface of the two-step coatings,as shown in Figure 9(b,c). This phenomenon is attributable tothe inhomogeneous thickness of the second PDLLA layer; theformation of HA was delayed in thick areas, relative to thinregions. Hence, the thickness of the second PDLLA coatingis an important factor in controlling both the mechanicaladherence and bioactivity of the coated sutures. A thickPDLLA film will certainly better protect Bioglass� particlesfrom peeling off than a thin one, at the same time it mayseverely retard the bioactivity of Bioglass� particles. Toachieve an optimal protocol, further investigation should fo-cus on a quantitative investigation involving a thickness-controllable coating technique, as well as understanding the

Figure 10. XRD patterns of two-step coated Vicryl� sutures after immersion in SBF for 4, 7, 14, and28 days, showing HA formation. The spectrum of an as-received Vicryl� suture is given for compar-ison. The coating slurry of the first step was 40 wt % Bioglass� in PDLLA–DMC solution. [Color figurecan be viewed in the online issue, which is available at www.interscience.wiley.com.]

361COMPOSITE COATINGS ON SURGICAL SUTURES

correlation between coating thickness, mechanical propertiesof the coated sutures, and their degradation rate.

A comparison between Figures 7 and 10 shows that theintensities of HA peaks are similarly strong for the one-stepcoating (Figure 7) and the two-step coating (Figure 10) at 4,7, 14, and 28 days of degradation in SBF. This result indi-cates that the second PDLLA thin film of the two-step coatingprocess did not block the bioactive response of the Bioglass�component of the coating. As mentioned above, this conclu-sion depends on the thickness of the second PDLLA layer.

As expected, uncoated sutures did not exhibit the forma-tion of apatite at all on their surfaces after immersion in SBFfor 4 days, as seen in Figure 11 for Vicryl� sutures. Table IIIpresents the main features of the coatings developed, sum-marizing the findings of this investigation.

The effect of time of immersion in SBF on the mechanicalproperties and degradation of PDLLA/Bioglass� coated suturesremains to be investigated. In our previous study,6 the tensilestrength of Bioglass� coated Vicryl� sutures, coated by thesimple slurry-dipping method, as a function of time of immer-

sion in SBF was investigated. It was found that the Bioglass�coating acts as a protective shield that affects both the extent andrate of degradation of the sutures. The rapid exchange of protonsin water for alkali in the glass provides a pH buffering effect atthe polymer surface. This behavior is also affected by the dis-solution of the glass leading to nucleation and growth of anhydroxyapatite layer as discussed above.

Our current work is focused on assessing the mechanicaland viscoelastic properties of the sutures coated by the two-step process and to quantify the variation of the suturesmechanical strength with degradation time in SBF. In partic-ular for Vicryl� sutures, it is expected that the addition of theBioglass� coating could be used to control the degradationkinetics of sutures.

CONCLUSIONS

A novel method has been developed for the coating of sur-gical sutures with bioactive glass particles and PDLLA films.Compared with pure Bioglass� coatings achieved by thesimple aqueous slurry dipping technique, the Bioglass�–PDLLA composite coatings exhibited improved microstruc-tural homogeneity and coating uniformity along the suturelength. By adding a second PDLLA coating, an increase ofthe mechanical adherence of the coatings to the sutures wasachieved. In vitro assessment in acellular SBF showed thatthe second PDLLA coating did not hindered the formation ofhydroxyapatite on the coated sutures, demonstrating theirbioactivity. These bioactive sutures, having a strongly ad-hered and bioactive polymer/Bioglass� coating, should beinteresting for the manufacture of scaffold architectures forbone tissue engineering using textile techniques.

We are grateful to Mr. Nick Royall and Mr. Matt Kershaw(Imperial College, London) for experimental assistance with SEMand XRD, respectively.

Figure 11. XRD patterns of uncoated Vicryl� sutures after immersion in SBF for 28 days and inas-received condition. There is no evidence of HA formation on the suture surface. [Color figure canbe viewed in the online issue, which is available at www.interscience.wiley.com.]

TABLE III. Features of the Bioglass�/PDLLA Coatings

1. HomogeneityBioglass� particles can be deposited homogeneously on suturesby the first step of the coating process.

2. Strong AdherenceThe mechanical stability of the coating can be achieved by thesecond PDLLA coating.

3. Good BioactivityThe thin PDLLA layer provides paths for the simulated bodyfluid to contact Bioglass� and stimulate bioactivity; meanwhileit maintains the mechanical integrity of the coating.

4. Key Factor of the Coating ProcessThe thickness of the second PDLLA layer is the significantfactor that controls the adherence and bioactivity of the finalcoating.

362 CHEN ET AL.

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