Full Terms & Conditions of access and use can be found athttp://www.tandfonline.com/action/journalInformation?journalCode=gpom20

Download by: [Tel Aviv University] Date: 03 May 2016, At: 03:29

International Journal of Polymeric Materials andPolymeric Biomaterials

ISSN: 0091-4037 (Print) 1563-535X (Online) Journal homepage: http://www.tandfonline.com/loi/gpom20

Novel Antibiotic-Eluting Gelatin-Alginate SoftTissue Adhesives for Various Wound ClosingApplications

Adaya Shefy-Peleg , Maytal Foox , Benny Cohen & Meital Zilberman

To cite this article: Adaya Shefy-Peleg , Maytal Foox , Benny Cohen & Meital Zilberman (2014)Novel Antibiotic-Eluting Gelatin-Alginate Soft Tissue Adhesives for Various Wound ClosingApplications, International Journal of Polymeric Materials and Polymeric Biomaterials, 63:14,699-707, DOI: 10.1080/00914037.2013.862535

To link to this article: http://dx.doi.org/10.1080/00914037.2013.862535

Accepted author version posted online: 16Apr 2014.

Submit your article to this journal

Article views: 164

View related articles

View Crossmark data

Citing articles: 1 View citing articles

Novel Antibiotic-Eluting Gelatin-Alginate Soft TissueAdhesives for Various Wound Closing Applications

ADAYA SHEFY-PELEG, MAYTAL FOOX, BENNY COHEN, and MEITAL ZILBERMAN

Department of Biomedical Engineering, Tel-Aviv University, Tel-Aviv, Israel

Received 19 September 2013, Accepted 22 October 2013

Interest in tissue adhesives as alternatives for conventional wound closing applications such as sutures and staples has increasedin the last few decades due to numerous possible advantages, including less discomfort and lower cost. In the current study,the authors developed novel tissue adhesives based on gelatin and alginate, crosslinked by carbodiimide. The antibiotic drugceftazidime was incorporated for prevention of infection. The bonding strength of the bioadhesives, their swelling ration,drug release profile and biocompatibility were studied. The high bonding strength and good biocompatibility turns these newbioadhesives into a promising alternative for use in wound closing applications.

Keywords: Alginate, bioadhesive, bonding strength, carbodiimide, ceftazidime hydrate, controlled drug delivery, cytotoxicity,gelatin

1. Introduction

Lacerations and traumatic wounds are considered to beamong the most prevalent scenarios treated in hospitalsand emergency rooms [1]. Re-attachment of the edges oflacerated tissues is traditionally carried out using suturesor staples. Use of tissue adhesives (i.e., substances that havethe ability to firmly attach lacerated tissues back together) asan alternative for these conventional applications has raisedinterest in the last few decades due to several major benefits[2,3]: Tissue adhesives can be applied more quickly, mayrequire less adhesive equipment, and are considered to bea relatively less time-consuming procedure. Use of tissueadhesives also prevents the painful procedure involved inusing sharp instruments and was proven to be less expensive,without compromising the cosmetic outcome [4]. Tissueadhesives can also be used to control bleeding, seal airleakage from the lungs, repair aortic dissections as well asfor external fixation of certain devices [5,6]. Tissue adhesiveshave a further potential for use as drug delivery systems.

An ideal soft tissue adhesive has not been developed todate despite extensive efforts that have been made in thepast, probably due to the various strict requirements thata substance must meet in order to serve as a tissue adhesivefor clinical use. The major requirements include efficientbonding strength to the tissue in a moist environment,sufficient biocompatibility to the tissue and its surrounding,

ease and convenience of application, sufficient flexibility forensuring that it will remain adhered to the tissue, penetrableto cell migration, biodegradability, stability during storage,and economical feasibility [7–9]. Nonetheless, a few productshave been approved for restricted medical use. These includecyanoacrylates, fibrin, and gelatin-based adhesives [5,6].

Cyanoacrylates create a strong bond to tissue, enablerapid hemostasis, and have the ability to polymerize uponcontact with fluids that are present on biological surfaces.However, they were found to be cytotoxic. Furthermore,the viscosity of non-cured adhesives is too low and curedadhesives are stiff and non-absorbable, and interfere withnormal wound healing [10]. Fibrin-based adhesives arehemostatic, biodegradable, promote wound healing, andadhere to connective tissues. Their disadvantages includelow strength (adhesive and cohesive) and low viscosity (hardto apply only to the desired site). They also pose a riskfor viral infection [10].

The main goal of the current study was to develop a noveltissue adhesive that meets all above-mentioned requirementsand to study the effects of its formulation parameters onthe resulting properties. Our new tissue adhesive is basedon the natural biopolymers gelatin and alginate, crosslinkedwith EDC (N-(3-dimethylaminopropyl)-N0-ethylcarbodiimidehydrochloride).

Gelatin is a natural water-soluble polymer derivedfrom collagen. It has become one of the most extensivelyinvestigated materials for tissue adhesives, due to its suitablenatural qualities. Gelatin is considered to be biocompatible,biodegradable and nonimmunogenic [11]. It can form physi-cally crosslinked hydrogel structures [12], has a natural tackybehavior in solution, and is very abundant in nature [13].

Address correspondence to: Prof. Meital Zilberman, Depart-ment of Biomedical Engineering, Faculty of Engineering, Tel-AvivUniversity, Tel-Aviv 69978, Israel. E-mail: meitalz@post.tau.ac.il

Color versions of one or more of the figures in the article canbe found online at www.tandfonline.com/gpom.

International Journal of Polymeric Materials and Polymeric Biomaterials, 63: 699–707

Copyright # 2014 Taylor & Francis Group, LLC

ISSN: 0091-4037 print/1563-535X online

DOI: 10.1080/00914037.2013.862535

Dow

nloa

ded

by [

Tel

Avi

v U

nive

rsity

] at

03:

29 0

3 M

ay 2

016

These characteristics have not only turned gelatin intoa promising candidate for tissue adhesives, but also fora wide range of other medical applications such as sealants[14], hydrogels [15], microspheres [16], and skin tissueengineering applications [17]. Despite its promising qualities,the mechanical strength of physically crosslinked gelatinadhesives is not sufficient as an adhering substance on itsown [13]. A chemical crosslinking agent and a polymericadditive (with appropriate available functional groups forthe crosslinking reaction) were therefore usually added tothe solution in a wide range of published attempts, in orderto create gelatin-based hydrogel formulations with suitablemechanical properties for soft tissue adhesion [8,13,18–20].

Carbodiimide, which is mainly used for modification andconjunction of proteins and other biological macrostructures,was chosen as the crosslinking agent since carbodiimides andtheir crosslinking byproducts have been reported to be lesscytotoxic than other conventional crosslinking agents suchas formaldehyde and glutaraldehyde [21].

Alginate is a natural polysaccharide that is extracted frommarine algae and is widely used in the food and beverageindustry as a gelling agent, stabilizer, and emulsifier [22]. Itis also applied in the medical and pharmaceutical industriesas a drug delivery vehicle [23–25], a dental impressionmaterial [26], part of a synthetic extracellular matrix for cellimmobilization [27] for wound dressing [28], and as filmsloaded with silver nanoparticle for inhibition of microbialgrowth in foodstuff [29]. Alginate was chosen as the poly-meric additive for the gelatin adhesive in the present study,as it is a natural source for high concentrations of carboxylicgroups, which are essential for the crosslinking reactionof carbodiimides. Carbodiimide binds to a carboxylic group(originally from the gelatin or the alginate) to form an o-iso-acylurea derivative, which is highly reactive and has anextremely short life. This activated structure goes througha nucleophilic attack by a primary amino group (originallyfrom the gelatin) to form an amide bond, and a ureamolecule (derivative of the carbodiimide type) is releasedas a byproduct [30]. As lacerated tissues contain exposedamino and carboxylic groups that can participate in thecrosslinking reaction, this adhesive has the potential forbeing especially attractive for tissue adherence applications.

Another new feature of our soft tissue adhesives, inaddition to the novel gelatin-alginate-carbodiimide combi-nation for bioadhesion applications, is the incorporation ofbioactive agents, such as antibiotic drugs, that are releasedto the surrounding wounded area and are therefore ben-eficial for prevention of infections. We chose to incorporateceftazidime pentahydrate in our adhesive. Ceftazidime isa third-generation cephalosporin, which has been shown tohave a broad spectrum of activity against Gram-positiveand Gram-negative bacteria, [31]. In one of our recentstudies we found controlled release of cetazidime frombiodegradable porous films and wound dressings derivedfrom freeze-dried inverted emulsions based on poly(DL-lactic–co–glycolic acid) [32,33]. The drug molecules werereleased in vitro within several days.

In the current study, the effects of the formulation para-meters on the adhesives’ function were studied. We examined

the effects of the gelatin, alginate, and the crosslinking agentconcentrations on the ability of the bioadhesives to bind tosoft tissues (an ex vivo model). The physical properties ofthe bioadhesives (i.e., swelling behavior and controlledantibiotic release, were also studied. The adhesives’ cyto-toxicity was evaluated on fibroblasts.

2. Experimental

2.1 Materials

Gelatin ‘‘type A’’ from porcine skin (90–100 bloom), alginicacid sodium salt (viscosity �250 cps, 2% [25�C]), N-(3-dimethylaminopropyl)-N’ ethylcarbodiimide hydrochloride(EDC), and ceftazidime hydrate (90–105%, C3809) were allpurchased from Sigma-Aldrich, Rehovot, Israel. The chemi-cal structure of the antibiotic drug ceftazidime hydrate is

2.2 Preparation of the Adhesive

Preparation of the adhesive was based on dissolving variousamounts of gelatin and alginate in double-distilled water,under heating up to 50�C. EDC was dissolved in double-distilled water and various amounts of this solution wereadded to the gelatin-alginate solution immediately priorto the adhesive’s use.

2.3 In Vitro Bonding Strength Measurements

Porcine skin (Kibbutz Lahav, Israel) was used as a softtissue model. The porcine skin was cut into 2� 2 cm2

square-shaped pieces and their epidermis side was firmlyattached to metal testing holders with a matchingsurface area (all dimensions of the holders are specified inFigure 1). 140 mL of the adhesive were then spread uniformlyon the dermis side of two porcine skin pieces (that wereattached to the testing holders) which were immediatelyattached by applying a 1.25N load on the pieces and placedin a 37�C and 100% humidity environment. After 30min,the bonding strength was measured in tension mode at roomtemperature using a 5500 Instron universal testing machine(Instron Engineering Corp.) and a 10N load cell. Thetwo parts of the joint were strained at a constant velocityof 2mm=min until separation was achieved. The mechanicaltesting procedure was inspired by the standard ASTMF-2258-03 test method. The bonding strength was calculatedbased on the average of five samples and was defined asthe maximum load divided by the bonding area. Additionalcuring times were also tested (1, 2, 5, and 10 h) aswell as the bonding strength of our adhesive with 1% w=vand 3% w=v ceftazidime. A one-way analysis of variancewith Tukey-Kramer multiple comparisons post-test wasused for group comparisons.

700 A. Shefy-Peleg et al.

Dow

nloa

ded

by [

Tel

Avi

v U

nive

rsity

] at

03:

29 0

3 M

ay 2

016

2.4 Swelling Ratio

The tissue adhesive was poured into a 6.2� 6.2� 3.5mm3

silicon mold and after gelation the adhesive was carefullyremoved and dried for 24 h. The adhesive was weighed(Wdry) and immersed in 2mL PBS (pH 7.0) with 0.02%w=v sodium azide as preservative. The sodium azide wasadded in order to prevent bacterial growth. The adhesiveswere placed in a static incubator at 37�C and 100% relativehumidity. The weight (Wwet) was measured after 3, 6, 9, 24,48, 72, and 96 h by removing the adhesive from the PBS andblotting using Kimwipes. The swelling ratio was calculatedaccording to Eq. 1:

Swelling ratio ð%Þ ¼ Wwet �Wdry

Wdry� 100 ð1Þ

2.5 In Vitro Drug Release Studies

The tissue adhesive was prepared in a 6.2� 6.2� 3.5mm3

silicon mold with 1% w=v and 3% w=v ceftazidime hydrate.After gelation the adhesive was carefully removed and driedfor 24 h. The adhesive was immersed in 2mL PBS with0.02% w=v sodium azide. The samples (n¼ 5) were immersedfor 14 days in order to determine the drug release kinetics.The adhesive samples were placed in a static incubator at37�C and 100% relative humidity. The medium was removedcompletely and fresh medium was added at 3, 6, 9, and 24 hand 2, 3, 7, and 14 days.

2.6 Residual Drug Recovery

The residual drug was extracted by dissolving the tissueadhesive in trypsin A for 24 h in a static incubator at 37�Cand 100% relative humidity. The residual drug was estimatedusing HPLC (as described subsequently).

2.7 Ceftazidime Assay

The medium’s ceftazidime content was determined usinga Jasco HPLC with a UV 2075 plus detector and a reversephase column (ACE 5 mm, inner diameter 4.6mm, length250mm), kept at 40�C. The mobile phase consisted ofa mixture of PBS (pH 5.0) and acetonitrile (94.5=5.5, v=v)at a flow rate of 1mL=min with a quaternary gradient pump(PU 2089 plus) without gradient. 20 mL samples wereinjected with an auto sampler (AS 2057 Plus). The columneffluent was eluted for 15min and detected at 254 nm.The area of each eluted peak was integrated using EZstartsoftware version 3.1.7. A calibration curve was preparedfor concentrations ranging from 5.0 to 400.0 mg=mL(correlation coefficient >0.9999).

2.8 Cytotoxicity Evaluation

The Alamar Blue assay was used to evaluate cell growth andviability in the presence of the bioadhesives. Human fibro-blasts were obtained from neonatal foreskins. Fibroblastcells (14th passage) were thawed and cultured in 75mm2

flasks with culture medium (modified Eagle’s medium[MEM] supplemented with heat-inactivated fetal bovineserum [FBS; 10% v=v], L-glutamine [1% v=v], and penicillin-streptomycin-nystatin [1% v=v]). The cultures were incubatedin a humidified atmosphere of 5% CO2 and 95% air. Thetissue culture medium was changed every 2 days until conflu-ence was reached. After the fibroblast confluence reached70% (14–20th passages), they were separated from thebottom of the flasks by a trypsin A solution and were seededinto 96-well plates at a concentration of 5000 cells per well.Former experiments (data not shown) determined thisoptimal seeding density and culture time. 0.2mL of freshculture medium were added to each well and the plates werereturned to incubation for 24 h until a confluence of 70% wasachieved in the wells.

In order to create the testing medium, 100mL ofgelatin-alginate solutions were mixed with 40mL EDC sol-ution with different concentrations (0–20mg=mL) to createcuboids. Every five cuboids were weighted, placed in scintil-lation vials and fresh culture medium was added at a ratio of1:5 (w=v). The scintillation vials were placed in a 37�C and5% CO2 humidified incubator environment for 24 h. TheAlamar Blue assay included replacing the original mediumwith 0.2mL of fresh medium containing 10% (v=v) AlamarBlue and incubation of the wells for 4 h. After the incu-bation, duplicates of 100mL from each well were transferredinto a 96-well plate for spectrophotometer analysis (Spectramax 340 PC384, Molecular Devices) at 570 and 600 nm. Thepercentage of Alamar Blue reduction was calculated accord-ing to the manufacturer’s protocol:

%AlamarBluereductionðEo

600ÞðA0570Þ� ðEo

570ÞðA0600Þ

ðEr570ÞðAc

600Þ� ðEr600ÞðAc

570Þ�100

ð2ÞWhere Eo and Errepresent the molar extinction coefficient

of the oxidized and the reduced Alamar Blue, respectively, at570 and 600 nm. A0 and Ao represent the absorbance of the

Fig. 1. Illustration of the bonding system for soft tissueadhesion measurements.

Antibiotic-Eluting Gelatin-Alginate Soft Tissue Adhesives 701

Dow

nloa

ded

by [

Tel

Avi

v U

nive

rsity

] at

03:

29 0

3 M

ay 2

016

test and control well (medium with Alamar Blue with nocells), respectively, at 570 and 600 nm. All samples weretested in triplicate.

3. Results

3.1 Bonding Strength

One of the main objectives in the current study was to achievehigh bonding strength, which is crucial in the application oftissue adhesives. We investigated the effect of various concen-trations of the adhesive components (gelatin, alginate, andEDC) on the bonding strength. A representative force-displacement curve presented in Figure 2 shows that afterfailure, the force slowly decreases, indicating that the failureis not brittle. However, the peak (maximal load) was usedfor calculating the bonding strength.

Figure 3a shows the effects of gelatin and alginateconcentrations on the bonding strength of the relevantformulations, crosslinked with EDC. It can be seen thatthere is a trend of higher bonding strength with theincrease in gelatin concentration. It is important to note thatalginate concentrations higher than 60mg=mL and gelatinconcentrations higher than 200mg=mL were highly viscousand therefore difficult to handle. Based on these results,a formulation of 200mg=mL gelatin and 40mg=mL alginatewas used for further studying the effects of the EDCconcentration on the bonding strength.

As expected, after the adhesive solution undergoesgelation it can no longer glue the skin. As a relatively highEDC concentration results in a short gelation time whichdoes not enable high quality bonding, a series of experimentswas conducted to examine the effect of the EDCconcentration (Figure 3b) using a slightly different appli-cation technique than in the series presented in Figure 3a.Figure 3b shows the effect of the EDC concentration onthe bonding strength of formulations, which combine200mg=mL gelatin and 40mg=mL alginate. As the EDCconcentration increases, there is a trend of increase inthe bonding strength. The maximal bonding strength wasmeasured for formulations crosslinked with 20mg=mLEDC. The gelatin (200mg=mL)-alginate (40mg=mL)-EDC(20mg=mL) formulation was therefore chosen for furtherstudy. EDC concentrations higher than 20mg=mL were

difficult to handle, due to the very fast gelation process. Fur-thermore, formulations with EDC concentrations of 30mg=mL resulted in a bonding strength of approximately6,300 Pa, which is lower than a similar formulation cross-linked with 20mg=mL EDC (approximately 10,000 Pa).The former also exhibited a higher standard error.

The effect of loading time on the bonding strength of thechosen formulation (gelatin [200mg=mL]-alginate [40mg=mL]-EDC [20mg=mL]) is presented in Figure 4. The

Fig. 2. Typical force versus displacement curve of the bioadhe-sives, shown for a formulation containing 200mg=mL gelatin,40mg=mL alginate, and 20mg=mL EDC.

Fig. 3. (a) The effect of gelatin and alginate concentrations onthe bonding strength of the tissue adhesives. Alginate concen-tration of 20mg=mL ( ), 40mg=mL ( ), and 60mg=mL ( ),EDC concentration 4mg=mL. (b) The effect of the EDC con-centration on the bonding strength of the tissue adhesive.Gelatin and alginate concentrations of 200mg=mL and 40mg=mL, respectively, were used. Values are expressed as mean�SE.

Fig. 4. The bonding strength of the chosen formulation (200mg=mL gelatin, 40mg=mL alginate, and 20mg=mL EDC) asa function of loading time. Values are expressed as mean�SD.

702 A. Shefy-Peleg et al.

Dow

nloa

ded

by [

Tel

Avi

v U

nive

rsity

] at

03:

29 0

3 M

ay 2

016

bonding strength remained practically unchanged for thefirst 5 h. However, a longer loading time resulted in anincrease in the bonding strength.

Additional formulations included the antibiotic drugceftazidime. As can be seen in Figure 5, penetration ofceftazidime molecules into the drug formulation results ina decrease in the bonding strength. The bonding strengthof the adhesive without the drug and the adhesive with 1%w=v ceftazidime was significantly higher than that of theadhesive with 3% w=v ceftazidime. The effect of 1% w=vdrug was not significant compared to the reference.

3.2 Swelling Ratio

Three types of adhesives were tested: adhesive withoutdrug, adhesive with 1% w=v ceftazidime, and adhesive with3% w=v ceftazidime. Time points of interest were those usedin the drug release experiment (see subsequent discussion).The adhesives were placed in PBS (pH 7, 37�C) and theswelling ratio was calculated according to equation 1 presentedin the methods section.

Figure 6 shows the swelling ratio of the chosenformulation adhesive at different times. The swelling ratioincreases with the increase in soaking time, until it reaches

a plateau. The swelling behavior can be divided into twostages:

a. A rapid initial water uptake within the first 3 h. Anincrease of 267%, 330%, and 368% w=w in the weight ofthe adhesive was observed at this stage for the chosenformulation without drug, adhesive with 1% w=v ceftazi-dime, and adhesive with 3% w=v ceftazidime, respectively.

b. A second stage consisting of a more moderate increase inthe water content during the following two days.

It is important to note that the swelling ratio increasedwith the increase in drug content.

3.3 In Vitro Ceftazidime Release Study

Drug release profiles of the chosen bioadhesive formulationloaded with 1% w=v and 3% w=v ceftazidime are presentedin Figure 7. These concentrations are equivalent to 3.7%w=w and 10.3% w=w (i.e., compared to the polymer(gelatin-alginate) mass. A burst release of 33% and 36%was obtained after 3 h for the 1% and 3% drug loads,respectively. After 24 h of release, the cumulative releasedceftazidime reached 45% and 47% of the encapsulated drugfor the 1% v=v and 3% v=v ceftazidime formulations,respectively. It remained the same also after 14 days (datanot shown). There was practically no difference betweenthe release kinetics of the two drug concentrations.

3.4 Cytotoxicity Evaluation

The cytotoxicity effect of the bioadhesives on fibroblast cellsis demonstrated in Figure 8. In general, a decrease in cell

Fig. 5. The effect of ceftazidime content on the bonding strengthof the chosen formulation (200mg=mL gelatin, 40mg=mLalginate, and 20mg=mL EDC). Values are expressed as mean�SEM. Significant differences are marked with an asterisk (�).

Fig. 6. Swelling ratio versus time of adhesives based on gelatin200mg=mL, alginate 40mg=mL, EDC 20mg=mL: without drug( ), 1% w=v ceftazidime ( ), and 3% w=v ceftazidime ( ).Values are expressed as mean�SD.

Fig. 7. In vitro release (mean�SD) of ceftazidime from the cho-sen bioadhesive formulation (gelatin 200mg=mL, alginate40mg=mL, EDC 20mg=mL) loaded with 1% w=v (^) and 3%w=v ( ) ceftazidime: (a) Cumulative release in %. (b) Cumulat-ive release in mg.

Antibiotic-Eluting Gelatin-Alginate Soft Tissue Adhesives 703

Dow

nloa

ded

by [

Tel

Avi

v U

nive

rsity

] at

03:

29 0

3 M

ay 2

016

viability results from an increase in the EDC concentration.At relatively low EDC concentrations (<15mg=mL), thecell viability is high (89–100%), while relatively high EDCconcentrations (15 and 20mg=mL) cause a decrease in cellviability, which is still higher than 70%. A similar behaviorwas also obtained after 48 h of incubation with the medium,which contained the bioadhesives.

Tissue adhesives based on gelatin and alginate crosslinkedwith EDC were developed and studied. Gelatin and alginatewere chosen due to their high biocompatibility, naturalorigin and low costs. EDC was chosen because it is reportedto have lower toxicity than other relevant crosslinkingagents, such as glutaraldehyde and formaldehyde [34].

3.5 Bonding Strength

Bonding of an adhesive to tissue starts with spreading theadhesive over the tissue in order to initiate close contact,which increases the contact surface area. Chains of theadhesive may diffuse into the tissue, thus increasing the con-tact area [35]. The bonding strength is the most importantproperty of tissue adhesives, whether there are used intern-ally, or externally for wound closure. The bonding abilityof gelatin-alginate solutions can be significantly increasedby the addition of a crosslinking agent.

Variability is common when testing adhesives in general,and even greater variability is expected when testing biologi-cal specimens. This variability is due to numerous factors:the tested tissue can have varying properties such as moist-ure levels and thickness, the alignment of the tissue priorto contact can be challenging, the adhesive joints can bedamaged in sample handling [36] and incomplete mixing ofthe tissue adhesive components or insoluble drug may causeinhomogeneous regions that can cause failure in some of thesamples [37]. The alignment device shown in Figure 1 wasdesigned to prevent alignment problems of the adhesivejoints, thus leading to more reproducible results.

As can be seen in Figure 3a, an increase in the gelatin con-centration increases the bonding strength due to more func-tional groups available for crosslinking. The same happens

when the alginate concentration is increased. Althoughgelatin contains both amino and carboxyl groups and canbe crosslinked by itself, the addition of alginate increasesthe concentration of carboxyl groups and should promotethe crosslinking with gelatin and reduce the gelation timeof the adhesive. However, the viscosity of an alginateconcentration of 60mg=mL is very high and is a possible causefor the low bonding strength at any gelatin concentration (i.e.,it is difficult to achieve a homogenous adhesive layer. Fora gelatin concentration of 100mg=mL, increasing the alginateconcentration did not increase the bonding strength becausethe gelatin concentration was relatively low. Low gelatinconcentrations weaken the anchoring effect that is caused bydeep infiltration of the adhesive into the tissue.

An increase in the EDC concentration results in highercrosslinking and therefore also in a higher bonding strength(Figure 3b). There is also a higher possibility for the forma-tion of amide bonds between gelatin and alginate, whichcontribute to higher bonding strength due to higher cohesionforces. However, there is probably an optimal EDC concen-tration (i.e., if it is too high, local crosslinking occurs instan-taneously and prevents the gelation of the rest of theadhesive, thus reducing the bonding strength. The increaseis not linear, probably due to microstructure reasons. It ispossible that crosslinking occurs in the range of 0–4mg=mL EDC, but not to an extent that significantly increasesthe bonding strength. The loading time effect on the bondingstrength was not significant (Figure 4). This shows that therewas probably no degradation in the humid environment dur-ing the first hours. The increase in the bonding strength after10 h is statistically significant. However, it may be caused bydehydration of the skin. These results show that the adhesivemaintains sufficient strength for at least 10 h. Further exam-ination for longer periods of time will be carried out in vivo.

The increase in drug content decreases the bondingstrength of the tissue adhesive as can be seen in Figure 5.These results can be explained by looking at the drug’schemical structure (see methods section). The drug’s primaryamine and carboxyl groups can be crosslinked with gelatinor alginate, reducing the crosslinking degree of the hydrogeland its bonding strength. The swelling results (Figure 6) thatshow a higher swelling ratio for bioadhesives loaded withceftazidime due to higher water uptake support this theory(i.e., the lower crosslinking density of the formulationsloaded with drug enables higher water uptake and swelling).

It is difficult to compare our bonding strength results toother studies on bioadhesives, due to differences in the testmethods (tensile, lap sheer or peel tests) and in curing con-ditions (temperature and % humidity), curing times andtissue types. For that reason we measured the bondingstrength of the fibrin sealant Evicel using our system andcompared the results with those measured for our formula-tions. The bonding strength of our chosen formulation(gelatin [200mg=mL]-alginate [40mg=mL]-EDC [20mg=mL]) is approximately 10,000 Pa, which is 4 times higherthan that of the fibrin sealant Evicel (approximately2500 Pa). Even when loaded with 3% w=v ceftazidime, ourchosen gelatin-alginate-EDC formulation enables muchhigher strength (approximately 5000 Pa). Furthermore, it

Fig. 8. Changes in fibroblast viability in the presence of thetissue adhesive (composed of 200mg=mL gelatin and 40mg=mL alginate) with various concentrations of EDC using theAlamar Blue assay, after incubation for 24 and 48 h. Valuesare expressed as mean�SD.

704 A. Shefy-Peleg et al.

Dow

nloa

ded

by [

Tel

Avi

v U

nive

rsity

] at

03:

29 0

3 M

ay 2

016

was difficult to handle the fibrin sealant because of the lowgelation time and low initial viscosity. The failure of thefibrin glue was adhesive in nature (failure which occursbetween the glue and the tissue), while the failure of ourtissue adhesives is usually cohesive, indicating betteradhesion to the tissue. It is possible that the gelatin-alginateadhesive is covalently linked to the tissue that containscarboxyl and amine groups.

3.6 Swelling Behavior

The degree of crosslinking is the most importantfactor affecting the swelling of biopolymers. Crosslinkinghinders the mobility of the polymer chain, thus loweringthe swelling ratio. Highly crosslinked biopolymers havea tighter structure and enable less swelling compared tobiopolymers with a lower degree of crosslinking [38].

The swelling ratio of a material can be used to estimate itscrosslinking degree. When comparing the swelling ratio ofan adhesive without drug with those of adhesives with 1%w=v and 3% w=v drug (Figure 6), it can be seen that theswelling ratio for each time point increases with the increasein drug content. This increase probably means that the cross-linking degree decreases. The decrease in crosslinking canexplain the reduction in bonding strength with the increasein drug content, as mentioned above. It is also possible thatthe hydrophilic clusters of ceftazidime absorb water.

3.7 Ceftazidime Release Study

The feasibility of the drug-eluting tissue adhesive was evalu-ated by the release of ceftazidime, and the results are pre-sented in Figure 7. The drug release mechanism in oursystem is probably diffusion, which is strongly enhancedby swelling of the gelatin-alginate matrix. Water penetratesthe adhesive, thus facilitating the diffusion of the drug. Asmentioned previously, the drug can be crosslinked with gela-tin or alginate due to the primary amine and carboxylgroups in its chemical structure. It is possible that this cross-linking of the drug molecules resulted in partial drug releaseof less than 50% of the encapsulated drug.

A local antibiotic release profile should exhibit a highinitial release rate in order to respond to the risk of infectionby bacteria during the initial shock of being wounded andkill bacteria that were not eliminated during cleansing [39].An additional sustained release phase at an effective anti-biotic level for inhibition is also desirable. Our profiles exhi-bit a good burst effect (33% and 36% for 1% w=v and 3%w=v ceftazidime, respectively). However, the duration ofrelease is relatively short (all drug release occurred within24 h). The burst release is more important than the prolongedrelease, which occurs later on, and we can therefore say thatfrom a practical point of view, the antibiotic’s release fromour bioadhesives is beneficial. Since the release rate ofwater-soluble drugs from biomaterials with medium or highwater uptake (hydrogels) is generally fast, immobilizationof the drug by electrostatic interactions between the drugand the releasing carrier may help achieve a prolonged releasetime. This could be carried out using a drug with a charge

opposite to that of gelatin. It is possible that a more hydro-phobic drug will also help extend the duration of drug release.

The rapid release of antibiotics from drug-eluting devicesis a well-known disadvantage that causes only a short-termantibacterial effect [40,41]. The reasons are usually thelow molecular weight and hydrophilic nature of the drugmolecules. Furthermore, the host polymeric matrix (inour case gelatin and alginate) is often hydrophilic, thusimposing little restriction on the drug’s diffusion. All ofthese reasons are also relevant in our drug-loaded bio-adhesives. The hydrophilic nature of ceftazidime enablesrelatively fast diffusion from the adhesive. The swellingresults may further explain the fast release, since thehydrogel rapidly absorbs water within the first 3 h. Therewas practically no difference in the release profiles ofthe two ceftazidime concentrations, probably due to thehydrophilic nature of the adhesive and the massive swellingduring the first hours.

3.8 Cytotoxicity

The Alamar Blue assay was performed on human fibroblaststhat participate in the wound healing process in order to testcell viability, and the entire experiment was performedaccording to a standard test method [42]. When tested separ-ately, the effect of the presence of gelatin or alginate on cellviability was similar to the control group [43] (i.e., gelatinand alginate are highly biocompatibile, as reported by otherresearches).

EDC is a coupling agent for amide bond formationbetween carboxyl and amino groups. The gelatin moleculescontain both types of groups, amino and carboxyl, andtherefore undergo inter and intra-molecular crosslinking byEDC. In this reaction, EDC is converted into a water-soluble urea derivative. Toxic effects may thus be obtaineddue to release of EDC remains to the surrounding wound,or due to the release of the urea derivative.

Figure 8 presents the cell viability of the bioadhesives asa function of the EDC concentration using the Alamar Blueassay. These results indicate that relatively low EDCcontents of 5 and 10mg=mL are preferable, as they are saferthan higher concentrations. Means for better mixing of EDCwith the polymeric medium may be used so as not to exceed10mg=mL EDC and still obtain a high bonding strength.These may spread the EDC more homogenously withinthe polymeric medium and therefore enable high strengthat low crosslinking agent contents. It is important to notethat a cell viability of more than 70% was achieved evenat relatively high EDC concentrations.

Other investigators who developed and studied bio-adhesives crosslinked with EDC also examined the cytotoxiceffects of EDC in their systems. For example, Otani et al. [8]reported that with a bioadhesive based on gelatin anda copolymer of lactic acid and glycolic acid, mice died whenthe EDC used in their bioadhesive was higher than 12.9mgin 270 mL implanted adhesive. Sekine et al. reported a tissueadhesive based on collagen and a copolymer of lactic acidand glycolic acid containing 12mg of EDC in 260 mLadhesive, that induced a more intense inflammatory reaction

Antibiotic-Eluting Gelatin-Alginate Soft Tissue Adhesives 705

Dow

nloa

ded

by [

Tel

Avi

v U

nive

rsity

] at

03:

29 0

3 M

ay 2

016

than fibrin glue in a rat model, but with no animal death[44]. Rahman et al. found that incorporation of urea ingelatin increased flexibility and stability in body fluid ofthe gelatin films and increases with no cytotoxic effect [45].

Our bioadhesive formulation with the maximal EDC con-centration contains approximately half of the EDC contentof the above-mentioned systems, and we can therefore say,taking this and the good cytotoxicity results into account,that our new bioadhesives are safe. It is also important tomention that the ceftazidime contents chosen in the currentstudy are expected to enable effective microbiological resultswithout inducing cytotoxic effects. This is based on oneof our previous studies on antibiotic release from wounddressings [46].

4. Conclusions

In the current study we developed and studied novelgelatin-alginate bioadhesives crosslinked with EDC andloaded with an antibiotic drug. Our thought was that inaddition to providing an attractive alternative for suturesand other traditional wound closing applications, ourbioadhesives will also prevent infection by local release ofceftazidime in the lacerated area. We first studied the effectsof the bioadhesives’ components on the bonding strengthusing an ex vivo porcine skin model and then used a selectedformulation for an advanced study, investigating the drugrelease profile and swelling behavior. The cytotoxicity ofthe bioadhesives was studied as well.

The bonding strength of the bioadhesives increased withthe increase in its component (gelatin, alginate, and EDC)concentrations. An optimal formulation was selected basedon the bonding strength results and practical aspects(viscosity, handling, etc.). The tissue adhesives maintaintheir bonding strength for at least 5 h in a humid environ-ment. Incorporation of drug (antibiotic) molecules decreasesthe bonding strength and increases the water uptake,probably due to a decrease in the crosslinking density.

The drug release experiment indicated burst release valuesof 33% and 36% during the first 3 h for 1% w=v and 3% w=vloaded ceftazidime, respectively, and the elution time lastedone day. The release mechanism is probably basedon diffusion enhanced by swelling of the gelatin-alginatematrix. The release rate of ceftazidime from the adhesive isrelatively high, due to the hydrophilic nature of both thedrug and the host polymer.

Finally, we demonstrated that proper selection of theadhesive component concentrations can yield antibiotic-eluting tissue adhesives with desired bonding strengths andphysical properties. Together with very high biocompatibi-lity at low-medium EDC contents, our current study maylead to tailoring of the bioadhesive for various medicalapplications.

Funding

The authors are grateful to the Office of the Chief Scientist(OCS) in the Israel Ministry of Industry, Trade and Labor,for supporting this research.

References1. Sibert, J. R.; Maddocks, G. B.; Brown, B. M. Arch. Dis. Childhood

1981, 56, 225–22.2. Osmond, M. H. Ped. Emerg. Care 1999, 15, 137–140.3. Sells, L.; Mihalov, L. Ped. Emerg. Med. Reports 1999, 4, 1–10.4. Osmond, M. H.; Klassen, T. P.; Quinn, J. V. J. Pediatr. 1995,

126, 892–895.5. Reece, T. B.; Maxey, T. S.; Kron, I. L. Am. J. Surg. 2001,

182, S40–S44.6. Ryou, M.; Thompson, C. C. Techn. Gastrointest. Endosc. 2006,

8, 33–37.7. Donkerwolcke, M.; Burny, F.; Muster, D. Biomaterials 1998,

19, 1461–1466.8. Otani, Y.; Tabata, Y.; Ikada, Y. J. Biomed. Mater. Res. 1996,

31, 157–166.9. Kraus, R.; Schluckebier, D.; Stiller, A. C.; Wenisch, S.; Schnettler,

R. Eur. J. Trauma 2006, 32, 141–148.10. Ratner, B. D.; Hoffman, A. S.; Schoen, F. J.; Lemons, J. E.

Biomaterials Science: An Introduction to Materials in Medicine;Elsevier Academic Press: San Diego, 2004.

11. Kuijpers, A. J.; Engbers, G. H. M.; Feijen, J.; De Smedt, S. C.;Meyvis, T. K. L.; Demeester, J.; Krijgsveld, J.; Zaat, S. A. J.;Dankert, J. Macromolecules 1999, 32, 3325–3333.

12. Djabourov, M.; Papon, P. Polymer 1983, 24, 537–542.13. Bae, S. K.; Sung, T. H.; Kim, J. D. J. Adhes. Sci. Technol. 2002, 16,

361–372.14. Jonas, R. A.; Ziemer, G.; Schoen, F. J.; Britton, L.; Castaneda,

A. R. J. Vasc. Surg. 1988, 7, 414–419.15. Tabata, Y.; Hijikata, S.; Ikada, Y. J. Control. Release 1994,

31, 189–199.16. Narayani, R.; Rao, K. P. J. Microencapsul. 1994, 11, 69–77.17. Choi, S. M.; Singh, D.; Kumar. A.; Oh, T. H.; Cho, Y. W.;

Han, S. S. Int. J. Polym. Mater. Polym. Biomater. 2013, 62,384–389.

18. Sung, H. W.; Huang, D. M.; Chang, W. H.; Huang, R. N.; Hsu,J. C. J. Biomed. Mater. Res. 1999, 46, 520–530.

19. Otani, Y.; Tabata, Y.; Ikada, Y. Biomaterials 1996, 17,1387–1391.

20. Matsuda, S.; Iwata, H.; Se, N.; Ikada, Y. J. Biomed. Mater. Res.1999, 45, 20–27.

21. Nishi, C.; Nakajima, N.; Ikada, Y. Wiley Online Library 1995,29, 829–834.

22. Draget, K. I.; Smidsrød, O.; Skjak-brak, G. Alginates from algae.Biopolymers 2005.

23. Liew, C. V.; Chan, L. W.; Ching, A. L.; Heng, P. W. S. Int. J.Pharmaceut. 2006, 309, 25–37.

24. Yang, H.; Wang, W.; Zhang, J.; Wang, A. Int. J. Polym. Mater.Polym. Biomater. 2013, 62, 369–376.

25. Mallikarjuna, B.; Madhusudana, R. K.; Siraj, S.; Chandra, B. A.;Chowdoji, R. K.; Subha, M. C. S. Des. Monomers Polym. 2013,16, 151–159.

26. Ashley, M.; McCullagh, A.; Sweet, C. Dental Update 2005,32, 169.

27. Smidsrød, O.; Skjaok-Brk, G. Trends Biotechnol. 1990, 8,71–78.

28. Matthew, I. R.; Browne S.; R. M.; Frame, J. W.; Millar, B. G.Biomaterials 1995, 16, 275–278.

29. Chaurasia, V.; Bajpai, S. Int. J. Polym. Mater. Polym. Biomater.2013, 62, 119–127.

30. Friess, W. Eur. J. Pharmaceut. Biopharmaceut. 1998, 45, 113–136.31. Richards, D. M.; Brogden, R. N. Drugs 1985, 29, 105–161.32. Rachelson, A.; Zilberman, M. Adv. Eng. Mater. 2009, 11.33. Elsner, J. J.; Zilberman, M. Acta Biomater. 2009, 5, 2872–2883.34. Sung, H. W.; Huang, D. M.; Chang, W. H.; Huang, R. N.; Hsu,

J. C. J. Biomed. Mater. Res. 1999, 46, 520–530.35. Lee, J. W.; Park, J. H.; Robinson, J. R. J. Pharmaceut. Sci. 2000,

89, 850–866.

706 A. Shefy-Peleg et al.

Dow

nloa

ded

by [

Tel

Avi

v U

nive

rsity

] at

03:

29 0

3 M

ay 2

016

36. McDermott, M. K.; Chen, T.; Williams, C. M.; Markley, K. M.;Payne, G. F. Biomacromolecules 2004, 5, 1270–1279.

37. Sierra, D. H.; O’Grady, K.; Toriumi, D. M.; Foresman, P. A.;Rodeheaver, G. T.; Eberhardt, A.; Lemons, J. E. J. Biomed. Mater.Res. 2000, 52, 534–542.

38. Peppas, N. A.; Hilt, J. Z.; Thomas, J. B. Horizon Biosci.2007, 10–11.

39. Zilberman, M.; Elsner, J. J. J. Control. Release 2008, 130, 202–215.40. Galdbart, J. O.; Branger, C.; Andreassian, B.; Lambert-Zechovsky,

N.; Kitzis. M. J. Surg. Res. 1996, 66, 174–178.41. Sripriya, R.; Kumar, M. S.; Sehgal, P. K. J. Biomed. Mater.

Res. Part B: Appl. Biomater. 2004, 70B, 389–396.

42. ISO 10993–12. Sample Preparation and Reference Materials;ISO-10993–5. Test for in Vitro Cytoxicity -(Indirect Contact).

43. Cohen, B.; Shefy-Peleg, A.; Zilberman, M. J. Biomater. Sci. Polym.Ed. 2014, 25, 224–240.

44. Sekine, T.; Nakamura, T.; Shimizu, Y.; Ueda, H.; Matsumoto,K.; Takimoto, Y.; Kiyotani, T. J. Biomed. Mater. Res. 2001,54, 305–310.

45. Rahman, M. F.; Islam, J. M. M.; Hassan, M. M.; Ahsan, H. S.M.; Pervez, M. S.; Khan, M. A. Int. J. Polym. Mater. Polym.Biomater. 2013, 62, 695–699.

46. Elsner, J. J.; Berdicevsky, I.; Zilberman, M. Acta Biomater. 2010,7, 325–336.

Antibiotic-Eluting Gelatin-Alginate Soft Tissue Adhesives 707

Dow

nloa

ded

by [

Tel

Avi

v U

nive

rsity

] at

03:

29 0

3 M

ay 2

016