http://www.diva-portal.org
This is the published version of a paper published in Advanced Healthcare Materials.
Citation for the original published paper (version of record):
Andren, O C., Ingverud, T., Hult, D., Håkansson, J., Bogestål, Y. et al. (2019)
Antibiotic-Free Cationic Dendritic Hydrogels as Surgical-Site-Infection-Inhibiting
Coatings
Advanced Healthcare Materials, : e1801619
https://doi.org/10.1002/adhm.201801619
Access to the published version may require subscription.
N.B. When citing this work, cite the original published paper.
Permanent link to this version:
www.advhealthmat.de
COMMUNICATION
Antibiotic-Free Cationic Dendritic Hydrogels as
Surgical-Site-Infection-Inhibiting Coatings
Oliver C. J. Andrén, Tobias Ingverud, Daniel Hult, Joakim Håkansson, Yalda Bogestål,
Josefin S. Caous, Kristina Blom, Yuning Zhang, Therese Andersson, Emma Pedersen,
Camilla Björn, Peter Löwenhielm, and Michael Malkoch*
DOI: 10.1002/adhm.201801619
(SSI).[1] The critical timeframe for
pre-venting an SSI is from approximately 1 h before the incision to 24 h after closure.[1,2]
A post-operative complication in the form of an SSI can negatively affect the healing process, resulting in an increase of the associated healthcare costs by approxi-mately one-third.[3] SSIs cause pain and
discomfort for patients and in 5% of cases mortality.[3] The current treatment
proto-cols can as an example consist of a single dose of antibiotic prophylaxis prior to inci-sion, with additional doses for lengthy pro-cedures.[4] Unfortunately, the widespread
consumption of antibiotics has led to the development of several multi drug-resistant (MDR) bacterial strains.[5] The inability
of current treatment protocols to address MDR bacteria is considered a major threat to public health.[6] We currently live in
an era where a simple bacterial infection can, once again, prove lethal.[7] One research area that has been
gaining traction for combating MDR is antimicrobial peptides (AMPs).[8] Unlike conventional antibiotics that mainly target the
internal components of a bacteria, AMPs mainly work through electrostatic forces damaging the bacterial membrane.[9]
How-ever, AMPs have limited use due to high production costs as well as toxicity toward mammalian cells.[10]
Peptide mimicking antimicrobial polymers have been pro-posed as an alternative to overcome the drawbacks associated with AMPs. With a similar mechanism of action and lower pro-duction price, antimicrobial polymers have shown a great poten-tial in combating the emergence of resistant bacteria.[11] SSIs
are generally local, rendering a systemic treatment excessive in most cases. One possibility of treating SSIs locally is through the use of an antimicrobial coating applied in the relevant sur-gical area.[12] There are several pathways to attain antimicrobial
properties in a coating, the two predominant ones capitalize on i) loading the material with a leachable antibiotic,[13] or ii) the
antimicrobial properties of the film forming components them-selves.[14] The former is a local drug delivery scaffold, while the
latter relies on antimicrobial compounds ranging from poly-mers to silver.[15] Most antimicrobial coatings presented to date
are hydrogels with the aim to coat implants or surfaces in a hos-pital setting but lack the possibility to coat tissue directly or con-tain components incompatible with an in vivo setting.[16]
A non-toxic hydrolytically fast-degradable antibacterial hydrogel is herein presented to preemptively treat surgical site infections during the first crucial 24 h period without relying on conventional antibiotics. The approach capitalizes on a two-component system that form antibacterial hydrogels within 1 min and consist of i) an amine functional linear-dendritic hybrid based on linear poly(ethylene glycol) and dendritic 2,2-bis(hydroxymethyl) propionic acid, and ii) a di-N-hydroxysuccinimide functional poly(ethylene glycol) cross-linker. Broad spectrum antibacterial effect is achieved by multivalent representation of catatonically charged β -alanine on the dendritic periphery of the linear dendritic component. The hydrogels can be applied readily in an in vivo setting using a two-component syringe delivery system and the mechanical properties can accurately be tuned in the range equivalent to fat tissue and cartilage (G′ = 0.5–8 kPa). The antibacterial effect is demonstrated both in vitro toward a range of relevant bacterial strains and in an in vivo mouse model of surgical site infection.
Dr. O. C. J. Andrén, T. Ingverud, Dr. D. Hult, Dr. Y. Zhang, Prof. M. Malkoch Division of Coating Technology
Department of Fibre and Polymer Technology School of Chemistry, Biotechnology and Health KTH Royal Institute of Technology
SE-100 44 Stockholm, Sweden E-mail: malkoch@kth.se T. Ingverud
Wallenberg Wood Science Center
Department of Fibre and Polymer Technology School of Chemistry, Biotechnology and Health KTH Royal Institute of Technology
SE-100 44 Stockholm, Sweden
Prof. J. Håkansson, Dr. Y. Bogestål, J. S. Caous, Dr. T. Andersson, Dr. E. Pedersen, Dr. C. Björn, Dr. P. Löwenhielm
RISE Research Institutes of Sweden Division Biosciences and Materials Section for Medical Device Technology Box 857, 50115 Borås, Sweden Dr. K. Blom
Medibiome AB 431 53 Mölndal, Sweden
The ORCID identification number(s) for the author(s) of this article can be found under https://doi.org/10.1002/adhm.201801619.
Surgical-Site Infections
Regardless of today’s surgical protocols, between 2% and 5% of all surgical interventions result in a surgical site infection
www.advancedsciencenews.com www.advhealthmat.de
A number of criteria needs to be fulfilled to truly utilize a hydrogel to prevent SSIs such as i) displaying a broad spectrum antimicrobial effect with pronounced efficacy against the most common bacterial strains including MDR strains, ii) ability to be formed in situ and in an in vivo environment with no harm done to surrounding tissue, iii) having sufficient mechanical proper-ties to ensure adequate coverage and contact to the intended tissue and avoid delamination, iv) maintaining the antimicrobial effect during the critical timeframe of post-surgery, and v) ide-ally be degradable with nontoxic degradation adducts.
Although antibacterial hydrogels have been studied extensively,[14,17] to the authors’ best knowledge, there are no
reports of in situ cured hydrogels designed for the specific purpose of protecting surgical sites the first 24 h followed by fast degradation to non-toxic constituents for preventing the onset of SSIs. In this work, a two-component hydrolytically degradable hydrogel system with tunable mechanical proper-ties was developed. The components of the system are i) An amine functional dendritic linear dendritic (DLD) hybrid pro-viding antibacterial properties and ii) a linear cross-linker with
N-hydroxysuccinimide (NHS) functionality.
The ester-based dendritic structure offers hydrolytic degra-dability as well as multivalence, while the PEG-based compo-nent offer solubility and biocompatibility. Drawing inspiration from nature, studying defense peptides, a scaffold of dense cationic nature was sought out. This was accomplished by
unfolding a library of DLDs being peripherally decorated with multitude of amino acid β-alanine. By varying the size of the dendritic block, from a 1st to a 6th generation, the amount of amines could be tuned from 4 up to 128 per molecule. All DLDs were synthesized from PEG-diol, molecular weights of 10 and 20 kg mol−1, with the monomer 2,2-bis(hydroxymethyl) propionic acid (bis-MPA) using divergent growth approach (Figure 1).[18] As shown in Figures S1–S7, Supporting
Informa-tion, the scaffolds were properly evaluated using 1H and 13C
NMR spectroscopy, MALDI-TOF-MS, and SEC. To introduce antibacterial properties, the scaffolds were converted to display amine functionality using tert-butyloxycarbonyl (BOC) pro-tected β-alanine followed by an acid deprotection step with trif-luoroacetic acid (TFA). The amine content was verified by poly electrolyte titration (PET) and the Kaiser test (see Table S1, Supporting Information).[19] To facilitate cross-linking in an
in vivo setting, three different telechelic NHS ester-based PEG cross-linkers, molecular weights of 2–10 kg mol−1, were syn-thesized (for additional details see Figure 1 and Table S1, Sup-porting Information).
Initially, the antibacterial properties of the amine func-tional PEG20k-based DLDs, from the 4th to the 6th generation containing from 32 up to 128 amines (38, 40, and 42), were investigated in solution using a minimal microbicidal concen-tration (MMC) assay, results are shown in Figure 2a. These DLDs were specifically selected due to their large number of
Figure 1. Synthetic scheme for both active amine functional dendritic linear dendritic (DLD)-hybrids (PEGy-Gx-(NH3+)2(x+1) where y is either 10 or
20 kg mol−1) and N-hydroxysuccinimide (NHS) PEG-based cross-linkers (PEGz-NHS where z is either 2, 6, or 10 kg mol−1). The DLD-hybrid represented here is just a 3rd generation scaffold.
www.advancedsciencenews.com www.advhealthmat.de
cationic charges and high aqueous solubility provided by the 20 kDa PEG chain. The MMC assay was performed using four different bacterial strains, including gram positive (S. aureus and P. aeruginosa) and gram negative (E. coli and E. faecalis). The antibacterial effect toward all strains was found to increase with increasing dendritic generation, with MMC values from approximately 3 µm for the 4th generation (38) down to less
than 1 µm for the 6th generation (42) scaffold (Figure 2a).
Increasing the amount of amines per molecule, influencing the density of amines represented by said molecule, appeared to have a direct effect on the antibacterial properties. The common antibiotic gentamicin sulfate was used as a positive control and had MMC values similar to the 4th generation scaf-fold. (Figure 2a).
The 6th generation scaffold (PEG20k-G6-(NH3+)128 (42)) was
due to its antibacterial effect selected for further studies and evaluated against E. faecealis, E. coli, P. aeruginosa, S. aureus, methicillin-resistant S. aureus (MRSA), S. epidermidis, and P. acnes in a minimal inhibitory concentration (MIC) assay. The material was found to have broad-spectrum effect, inhibiting growth of tested bacteria in concentrations down to the single digit nano-molar range (Figure 2a). Furthermore, the cytotoxicity of 42 was evaluated with human dermal fibroblasts (hDF) using an AlamarBlue assay. No cytotoxicity was observed up to 10 µm or
slightly above 0.1 wt% (Figure 2b), defined as above 70% viability. For the 6th generation scaffold broad-spectrum antibacterial
activity was observed at around 1 µm in MMC and below 1 nm in
MIC, in other words, well below the limit of cytotoxicity. The antibacterial mechanism of the polymer was investi-gated with both E. coli and S. Aureus using two methods, that is, i) bacteria were loaded with the self-quenching fluorescent dye disC3-5 and exposed to the test substance,[20] and ii)
anti-bacterial DLD was fluorescently tagged and co-localized in bac-teria discerning where it was localized in both dead and alive bacteria. Both S. aureus and E. coli loaded with fluorescent dye and exposed to PEG20k-G6-(NH3+)128 (42) displayed release of
the loaded fluoresce, indicating that the bacterial membrane was ruptured (Figure 2d and Figure S8, Supporting Informa-tion). In addition, pacific blue (PB) labeled PEG20k-G6-(PB)2
-(NH3+)126 (43) was utilized to fluorescently trace the polymer
in S. Aureus and E. coli. After 45 min exposure, alive bacteria (green stain) displayed a halo of polymer (blue florescence) indicating that the polymer was localized in the membrane. Dead bacteria (red stain) had blue signal through the bacteria indicating that the membrane was ruptured (see Figure 2e,f). The results for S. Aureus were similar to those for E. coli and can be found in Supporting Information along with additional details (Figures S8 and S9, Supporting Information). Both methods indicate that the antibacterial effect is exerted through membrane penetration. Satisfied with the antibacterial and bio-logical results based on PEG20k-G6-(NH3+)128 (42), hydrogels
were developed by statistical cross-linking at different molar
Figure 2. Evaluation of amine functional DLDs antibacterial properties in solution. a) Minimal microbicidal concentration (MMC) assay (µmolar polymer) for PEG20k-G4-(NH3+)32 (38), PEG20k-G5-(NH3+)64 (40), PEG20k-G6-(NH3+)128 (42) using gentamicin sulfate as positive control. Minimal
inhibitory concentration (MIC) for 42 (nmolar polymer). Bacterial strains where used: S. aureus, P. aeruginosa, E. coli, and E. faecalis, methicillin-resistant
S. aureus (MRSA), S. epidermidis and P. acnes. ND, not determined. b) Evaluation of cytotoxicity of 42 in human dermal fibroblasts (hDF) and Raw
264.7 cell lines using a AlamarBlue assay; data are presented as a mean value (n = 5). c) Loaded self-quenching fluorescent release from E. coli upon addition of 42; data are presented as a mean value (n = 5). d) Localization of PEG20k-G6-(PB)2-(NH3+)126 (43) (blue and pink) in E. coli together with
www.advancedsciencenews.com www.advhealthmat.de
ratios of peripheral amines, maintaining a balance between preserved amines with cationic features and amines available for cross-linking. To optimize the two-component formulation for use in a dual syringe delivery system (Figure 3a), the in situ cross-linking kinetics were studied in a rheometer. Hydrogels with inherently different concentration of cationic charges were constructed through statistical cross-linking using 4, 8, 16, and 32 (n) amines out of the total 128 for PEG20k-G6-(NH3+)128 (42)
together with the telechelic PEG10K-NHS (10) cross-linker. All hydrogels maintained a solid content of 10 wt%. The effect of number of cross-links per DLD were evaluated in terms of mechanical properties and cross-linking time as illustrated in Figure 3b.
In situ curing experiments were carried out in a rheometer to determine storage modulus (G′) and cross-linking time. As shown in Figure 3e and Figures S10–S14, Supporting Informa-tion, the results revealed that four cross-links per molecule affords a G′ of 0.5 kPa, equivalent to brain and fat tissue.[21] Increasing
the amount of cross-links from four to eight resulted in a substan-tial increase in G′, reaching 2.4 kPa in the upper modulus range of skin or the lower range of muscle.[21] Further increasing the
number of cross-links to 16 and 32 yielded moderate gains in G′ up to 3.7 kPa and 4.9 kPa, respectively. These values are equivalent to the higher range of muscle or lower range of tendon tissue.[21]
Consequently, a cross-linking density of 8 out of the 128 amines was recognized as optimal in terms of modulus gain (close to skin and muscle as our targeted tissue) and maximal availability of cationic amines, for antibacterial effect. It should be noted,
materials with modulus close to the targeted tissue ensures com-patibility with surrounding tissue which reduce scar-formation[22]
as well as promote cell differentiation.[23]
In the case of time-dependent cross-linking, the goal was to identify a hydrogel system that exhibit sufficiently slow kinetics for feasible application, while retaining rapid enough curing suitable for surgical use. In this context, a cross-linking time of around 1 min was considered optimal. Opting for 8 rather than 4 theoretical cross-links noted faster curing cycle of the hydrogels with a drop from 87 down to 54 s. Increasing the number of cross-links beyond 8 came with modest improve-ments to curing kinetics having 49 and 48 s for 16 and 32, respectively.
The degradation profile of hydrogels is a crucial prop-erty to assess. Therefore, the degradation kinetics of the hydrogel system based on PEG20k-G6-(NH3+)128 x8
PEG10k-NHS was investigated mechanically. Samples cured in situ a rheometer were submerged in aqueous buffers of pH 7, 8, 9, and 10 and the modulus was monitored in real time. Furthermore, human blood plasma was used to distin-guish between the hydrolytic and the enzymatic contri-bution to degradation (see Figure 3d). The initial slope of the degradation profiles was used to determine degra-dation rate in Pa s−1. Results indicate that the degrada-tion of ester bonds is primarily of hydrolytic nature rather than enzymatic (see Figure 3g and Figure S17, Supporting Information). In human blood plasma at pH 7.4, the gels were degraded below a modulus of 0.1 kPa after 22.5 h.
Figure 3. Cross-linking evaluation using amine functional DLDs and NHS functional telechelic PEGs as cross-linkers. a) Dual syringe application
system with air assisted spray attachment. b) Illustration of varying degrees of cross-linking. c) Illustration of varying lengths of linear segments for both the amine functional DLD hybrid and the NHS functional PEG cross-linker. d) Normalized storage modulus (G′/G′0) as a function of time
during degradation at four different pH (7, 8, 9, and 10) and in blood plasma as measured by a time sweep in the rheometer with the sample submerged in the medium at 37 °C for PEG20k-G6-(NH3+)128 x8 PEG10k-NHS. e) Time of cross-linking and modulus as a function of varying degrees
of theoretical cross-linking for PEG20k-G6-(NH3+)128 x (n) PEG10k-NHS where n is 4, 8, 16, or 32. f) Theoretical molecular weight between cross-links,
theoretical mm amine m−3 and measured modulus as a function of varying the length of the linear block in both amine functional DLD and the NHS
functional cross-linker (PEGy-G6-(NH3+)128 x8 PEGz-NHS where y is either 10 or 20 kg mol−1 and z is 2, 6, and 10 kg mol−1); data are presented as a
mean value (n = 3). g) Linear slope extracted from the initial segment of the degradation time sweep plotted against the pH with confidence bands and a linear curve fit.
www.advancedsciencenews.com www.advhealthmat.de
Therefore, the degradation time of the hydrogels appears to only depend on the initial modulus and the pH of the sur-rounding medium.
In-depth assessment of the degradation profile for the cross-linker PEG10k-NHS, with labile succinic ester bonds, was per-formed by 1H-NMR spectroscopy. Approximately one-fourth of
the esters were found to hydrolyze rapidly at pH 7.4 after which degradation was considerably slower as seen in Figure S18, Supporting Information. The retardation observed is probably due to acidification generated by increased number of carbox-ylic end groups. This requires that dissolution and mixing of cross-linker and DLDs to be rapid, something handled well by the dual syringe system used in this study. The swelling of the hydrogel was also investigated and compared to the degrada-tion profile. In this case, the hydrogels were found to swell approximately 200 +% based on dry weight and follow expected profile up to about 3 h; data can be seen in Figure S19, Sup-porting Information.
To further understand the composition impact on hydrogel modulus, the average number of cross-links per molecule was maintained at eight and the length of the linear PEG seg-ments (y and z) was varied, both directly influencing the length between cross-links as illustrated in Figure 3c. By varying the length of the linear segment for the amine functional DLD (PEGy-G6-(NH3+)128 where y is either 20 or 10 kg mol−1) and
by varying the length of the cross-linker (PEGz-NHS where z is either 2, 6, or 10 kg mol−1) the overall density and properties of the networks could be altered. We postulated that a denser net-work, that is, with a shorter distance between amine-clusters, would result in a higher antibacterial effect. This was based on the observed increased antibacterial effect with increase amount of amines in solution. The density of the network should also be proportional to the final modulus. The results can be seen in Figure 3f together with a theoretical average molecular weight between cross-links and amount of amines per m3. An increase
of modulus from approximately 6.5 up to 8.0 kPa was observed when the PEG length of the linear segment in the amine func-tional DLD was decreased from 20 to 10 kDa. However, when the length of the cross-linker was decreased, 6 and 2 kDa, a decrease in modulus was observed. This could possibly be due to the formation intramolecular loops instead of intermolecular cross-links caused by the decreased distance between chain ends due to decreased molecular weight. Collectively, the mod-ularity presented with this approach can indeed be translated to the development of antibacterial hydrogels with controlled modulus and degradation profile, both of high importance for medical devices.[24]
The formulation containing the shortest linear segments (PEG10k-G6-(NH3+)128 x8 PEG2k-NHS) was discarded since the
modulus was too low for adequate physical stability while all other systems, shown in Figure 3f, were evaluated in the dual syringe delivery system and where found to form soft coat-ings. All three formulations were further analyzed with regard to their antibacterial effect in a modified MMC gel assay, using polyhexamethylene biguanide (PHMB) in solution as a control (Figure 4c). A 24-wellplate with horse blood agar was used where each well was coated with individual hydrogel for-mulation. Various hydrogel thicknesses of 25, 50, 100, 200, and 400 µm were introduced to investigate their antibacterial
effect. The bacteria (S. aureus) were then deposited on top of the hydrogel and the growth counted as colony-forming units (CFU) after 16 h. The survivability was compared against the amount added CFU per well and the most stable results here obtained for between 25 and 250 CFU (see Table S3, Supporting Information). Hydrogels with thickness of 400 µm exhibited the strongest antibacterial effect. At the optimal combination of gel thickness and fixed CFU of 100 per well the gels were then further analyzed against MRSA, E. coli and P. aeruginosa. Notably for MRSA, no bacterial colonies were formed (see Figure 4c). These results are very promising since MRSA is of critical concern when discussing SSIs.[25] In both E. coli
and P. aeruginosa, the highest effect could be observed for the longest distance between cross-links (PEG20k-G6-(NH3+)128 x8
PEG10k-NHS) and appeared to decrease with decreasing length between cross-links, being the lowest for PEG10k-G6-(NH3+)128
x8 PEG6k-NHS.
In the case of S. aureus, a trend similar to E. coli and
P. aeruginosa was observed, were the lowest effect was observed
for the shortest cross-linking length. The PEG20k-G6-(NH3+)128
x8 PEG10k-NHS system had over all the highest antibacte-rial effect across all strains with an average of 88% reduction, closely followed by 85.5% for the (PEG10k-G6-(NH3+)128 x8
PEG10k-NHS) system and 76.8% for the (PEG10k-G6-(NH3+)128
x8 PEG6k-NHS) system.
The biocompatibility of the hydrogels and their degradation adducts were also assessed by a minimum essential media (MEM) elution test using hDF and RAW cell-lines. Overall results detailed that the hydrogels and degradation adducts were non-cytotoxic at all time-points. The time-points were selected to feature varying degrees of degradation and is correlated with the degradation profile of the hydrogels from Figure 3(d and g), (see Figure 4a and the RAW results in Figure S20, Supporting Information).
Finally, the most promising two component (PEG20k-G6-(NH3+)128 x8 PEG10k-NHS) hydrogel system was used with
a dual syringe (illustrated in Figure 3a) for an easy on-surgical-site application in an in vivo animal mouse model of a harsh simulated SSI using an S. aureus infected suture (Figure 4b), as a proof-of-concept. The clinical indication for this product is a surgical site with a low infection dose (few CFU). This is difficult to mimic in an animal model. Instead, we use an established in vivo model with an infection load of 106 CFU
to obtain a statistically more stable model and thereof prop-erly study the antibacterial effect. Despite the high infection dose, 80% reduction of bacteria was observed 4 h post-infec-tion and a full log (90%) reducpost-infec-tion after 6 h post-infecpost-infec-tion in vivo (Figure 4d). The formulations ability to handle the harsh conditions with an initial infection inoculum of 106 CFU
strengthens our initial hypothesis that the antibacterial den-dritic hydrogels could be utilized to inhibit the initial infection from taking grip.
To summarize, DLD-hybrids based on linear PEG and bis-MPA dendritic polyesters, from the 1st to the 6th generation, were synthesized and modified to display multitude of cati-onic amines. The hybrids antibacterial effect was investigated in vitro and the-6th generation scaffolds were found to have a broad spectrum antibacterial effect. The observed antibacte-rial effect increased with increasing dendritic generation and
www.advancedsciencenews.com www.advhealthmat.de
the highest effect was observed for the 6th generation amine functional DLD-hybrid. The 6th generation DLD-hybrid was further screened toward a spread of bacterial strains commonly found in SSIs and displayed antibacterial effect toward all strains tested. Hydrogel formulations was optimized with respect to structural compositions and under physiological con-ditions using NHS functional linear PEG as cross-linkers. The hydrogels were found to undergo hydrolytic degradation and the mechanical properties where tunable to cover the range of important storage modulus of soft tissues, ranging from 0.5 to 8 kPa. A dual syringe delivery system developed and validated, preclinical on a mouse model, for the formation of hydrogels within 1 min after application. The hydrogels antibacterial and growth inhibitory effect were evaluated in vitro with broad spectrum antibacterial effect, and in vivo with a full log reduc-tion after 6 h in a severe infecreduc-tion model. To conclude a new approach and method to proactively treat SSIs was developed relying on quick cross-linking antibacterial polymer networks made possible with dual-syringe application under a surgical setting.
Supporting Information
Supporting Information is available from the Wiley Online Library or from the author.
Acknowledgements
O.C.J.A. and T.I. contribute equally to this work. O.C.J.A., M.M. and P.L. conceptualized the work. O.C.J.A. and T.I. preformed the synthesis. O.C.J.A., T.I., and D.H. made formulations and analyzed the polymers and gels. J.H. and Y.B. conceptualized and performed the animal study. J.S.C., K.B., Y.Z., T.A., E.P., and C.B. performed the microbial and cytotoxic analysis. O.C.J.A., M.M., P.L., T.I., D.H., and J.H. wrote the manuscript. The authors acknowledge Knut and Alice Wallenberg Foundation (KAW), Seventh Framework Program and Swedish Research Council (VR) for financial support. All animal experiments were performed after prior approval from the local ethics committee for animal studies at the administrative court of appeals in Gothenburg, Sweden.
Conflict of Interest
The authors declare no conflict of interest.
Figure 4. Effect evaluation of antibacterial hydrogels. a) MEM elution cytotoxicity test using hDF cells, was used as is or diluted 10 and 100 times.
The weight in the legend refers to the total amount of hydrogel added to the media. b) Illustration of animal model. c) Modified minimal microbicidal concentration (MMC) gel assay preformed on gel surfaces testing survivability of S. aureus, MRSA, E. coli, and P. aeruginosa. Gels were formed from PEG20k-G6-(NH3+)128 x8 PEG10k-NHS, PEG10k-G6-(NH3+)128 x8 PEG10k-NHS and PEG10k-G6-(NH3+)128 x8 PEG6k-NHS (160 µL total of each gel with
a dry content of 10 wt% was applied to each well). Initial bacterial dose 100 CFU; data are presented as a mean value (n = 5). d) Animal study using infected suture (106 CFU) model. The graph contains both bacterial reduction in percent of control and log CFU based reduction of using
antibacte-rial hydrogel. Each wound was treated with 100 µL of PEG20k-G6-(NH3+)128 x8 PEG10k-NHS gel with a dry content of 10 wt% (Mann–Whitney U test.
www.advancedsciencenews.com www.advhealthmat.de
Keywords
antibacterial, dendrimer, hydrogels, surgical-site infection
Received: December 18, 2018 Revised: January 19, 2019 Published online: February 8, 2019
[1] D. J. Anderson, K. Podgorny, S. I. Berrios-Torres, D. W. Bratzler, E. P. Dellinger, L. Greene, A. C. Nyquist, L. Saiman, D. S. Yokoe, L. L. Maragakis, K. S. Kaye, Infect. Control Hosp. Epidemiol. 2014, 35, 605.
[2] a) M. McDonald, E. Grabsch, C. Marshall, A. Forbes, Aust. N. Z. J.
Surg. 1998, 68, 388; b) B. A. Coakley, E. S. Sussman, T. S. Wolfson,
A. S. Bhagavath, J. J. Choi, N. E. Ranasinghe, E. T. Lynn, C. M. Divino, J. Am. Coll. Surg. 2011, 213, 778.
[3] a) R. E. Nelson, M. L. Schweizer, E. N. Perencevich, S. D. Nelson, K. Khader, H. Y. Chiang, M. L. Chorazy, A. Blevins, M. A. Ward, M. H. Samore, Infect. Control Hosp. Epidemiol. 2016, 37, 1212; b) E. Zimlichman, D. Henderson, O. Tamir, C. Franz, P. Song, C. K. Yamin, C. Keohane, C. R. Denham, D. W. Bates, JAMA Intern.
Med. 2013, 173, 2039.
[4] a) R. L. Nichols, Clin. Med. Res. 2004, 2, 115; b) T. Hranjec, B. R. Swenson, R. G. Sawyer, Surg. Infect. 2010, 11, 289.
[5] S. B. Levy, B. Marshall, Nat. Med. 2004, 10, S122.
[6] a) L. B. Rice, J. Infect. Dis. 2008, 197, 1079; b) L. B. Rice, Infect.
Control Hosp. Epidemiol. 2010, 31, S7.
[7] a) T. U. Berendonk, C. M. Manaia, C. Merlin, D. Fatta-Kassinos, E. Cytryn, F. Walsh, H. Burgmann, H. Sorum, M. Norstrom, M.-N. Pons, N. Kreuzinger, P. Huovinen, S. Stefani, T. Schwartz, V. Kisand, F. Baquero, J. L. Martinez, Nat. Rev. Microbiol. 2015,
13, 310; b) E. D. Brown, G. D. Wright, Nature 2016, 529, 336;
c) Antimicrobial Resistance: 2014 Global Report on Surveillance, World Health Organization, 2014.
[8] K. A. Brogden, Nat. Rev. Microbiol. 2005, 3, 238.
[9] a) M. Zasloff, Nature 2002, 415, 389; b) M. A. Kohanski, D. J. Dwyer, J. J. Collins, Nat. Rev. Microbiol. 2010, 8, 423.
[10] R. E. W. Hancock, H. G. Sahl, Nat. Biotechnol. 2006, 24, 1551. [11] a) M. F. Ilker, K. Nusslein, G. N. Tew, E. B. Coughlin, J. Am. Chem.
Soc. 2004, 126, 15870; b) K. Kuroda, W. F. DeGrado, J. Am. Chem. Soc. 2005, 127, 4128; c) G. N. Tew, D. Clements, H. Z. Tang,
L. Arnt, R. W. Scott, Biochim. Biophys. Acta, Biomembr. 2006, 1758, 1387; d) G. J. Gabriel, A. Som, A. E. Madkour, T. Eren, G. N. Tew,
Mater. Sci. Eng., R 2007, 57, 28; e) E. R. Kenawy, S. D. Worley,
R. Broughton, Biomacromolecules 2007, 8, 1359.
[12] a) D. Campoccia, L. Montanaro, C. R. Arciola, Biomaterials 2013,
34, 8533; b) F. Siedenbiedel, J. C. Tiller, Polymers 2012, 4, 46.
[13] K. Malizos, M. Blauth, A. Danita, N. Capuano, R. Mezzoprete, N. Logoluso, L. Drago, C. L. Romano, J. Orthop. Traumatol. 2017,
18, 159.
[14] P. Li, Y. F. Poon, W. F. Li, H. Y. Zhu, S. H. Yeap, Y. Cao, X. B. Qi, C. C. Zhou, M. Lamrani, R. W. Beuerman, E. T. Kang, Y. G. Mu, C. M. Li, M. W. Chang, S. S. J. Leong, M. B. Chan-Park, Nat. Mater.
2011, 10, 149.
[15] a) L. Ferreira, A. Zumbuehl, J. Mater. Chem. 2009, 19, 7796; b) J. Lin, S. Y. Qiu, K. Lewis, A. M. Klibanov, Biotechnol. Prog. 2002,
18, 1082; c) J. C. Tiller, C. J. Liao, K. Lewis, A. M. Klibanov, Proc. Natl. Acad. Sci. USA 2001, 98, 5981; d) J. Henschen, P. A. Larsson,
J. Illergard, M. Ek, L. Wagberg, Colloids Surf., B 2017, 151, 224; e) C. Chen, J. Illergard, L. Wagberg, M. Ek, Holzforschung 2017, 71, 649; f) J. Illergard, L. Wagberg, M. Ek, Cellulose 2015, 22, 2023. [16] a) M. J. Li, L. Y. Gao, C. Schlaich, J. G. Zhang, I. S. Donskyi,
G. Z. Yu, W. Z. Li, Z. X. Tu, J. Rolff, T. Schwerdtle, R. Haag, N. Ma,
ACS Appl. Mater. Interfaces 2017, 9, 35411; b) H. Han, J. F. Wu,
C. W. Avery, M. Mizutani, X. M. Jiang, M. Kamigaito, Z. Chen, C. W. Xi, K. Kuroda, Langmuir 2011, 27, 4010.
[17] a) L. Shuqiang, D. Shujun, X. Weiguo, T. Shicheng, Y. Lesan, Z. Changwen, D. Jianxun, C. Xuesi, Adv. Sci. 2018, 5, 1700527; b) Y. Luo, K. R. Kirker, G. D. Prestwich, J. Controlled Release 2000,
69, 169.
[18] a) G. R. Newkome, Z. Q. Yao, G. R. Baker, V. K. Gupta, J. Org. Chem.
1985, 50, 2003; b) D. A. Tomalia, H. Baker, J. Dewald, M. Hall,
G. Kallos, S. Martin, J. Roeck, J. Ryder, P. Smith, Polym. J. 1985,
17, 117; c) E. Buhleier, W. Wehner, F. Vogtle, Synthesis 1978, 1978,
155.
[19] E. Kaiser, R. L. Colescott, C. D. Bossinger, P. I. Cook, Anal. Biochem.
1970, 34, 595.
[20] T. A. Chitemerere, S. Mukanganyama, BMC Complementary Altern.
Med. 2014, 14.
[21] S. Sartori, V. Chiono, C. Tonda-Turo, C. Mattu, C. Gianluca, J. Mater.
Chem. B 2014, 2, 5128.
[22] K. C. Spencer, J. C. Sy, K. B. Ramadi, A. M. Graybiel, R. Langer, M. J. Cima, Sci. Rep. 2017, 7.
[23] a) R. J. Wade, J. A. Burdick, Mater. Today 2012, 15, 454; b) Z. Z. Khaing, R. C. Thomas, S. A. Geissler, C. E. Schmidt, Mater.
Today 2014, 17, 332.
[24] R. P. Brannigan, A. P. Dove, Biomater. Sci. 2017, 5, 9.
[25] a) F. A. Manian, P. L. Meyer, J. Setzer, D. Senkel, Clin. Infect.
Dis. 2003, 36, 863; b) R. A. Zoumalan, D. B. Rosenberg, Arch. Facial Plast. Surg. 2008, 10, 116; c) K. S. Kaye, D. J. Anderson,
Y. Choi, K. Link, P. Thacker, D. J. Sexton, Clin. Infect. Dis. 2008, 46, 1568.
