Patent Application
20250121121
The present disclosure relates generally to surface treatments, and more particularly to methods for imparting resistance to bacterial colonization to substrates through treatment with peptoid-loaded microgels.
Even though surgical operating rooms (ORs) are commonly referred to as being sterile or aseptic, in reality, OR atmospheres contain microbes from many different sources. Such sources include ventilation systems, shedding from clothing, sneezing or coughing by OR personnel, and pedestrian traffic [1]. These airborne bacteria can sediment directly onto the surfaces of an implantable biomedical device once the device is removed from its sterile packaging. Early measurements [2] reported a sedimentation rate of about 3×106 CFU/m2-h. Improved sanitation practices have steadily reduced sedimentation rates to about 103 to 104 CFU/m2-h [3]. However, these sedimentation rates suggest that, even under these improved conditions, devices are often being implanted after they have been contaminated by hundreds or thousands of bacteria. Once implanted and exposed to favorable growth conditions within the body, a subset of these bacteria may develop into biofilms and can lead to chronic, device-associated infections.
Various implantable devices, such as hip/knee prostheses [4], heart valves [5], pacemakers [6], cochlear implants [7], shunts [8], surgical mesh [9], sutures [10], and tissue-engineering constructs [11], are susceptible to device-associated infection. It is well established that the incidence of surgical site infection increases linearly with time in the OR [12], and there is a strong consensus that intra-operative contamination is responsible for at least some device-associated infections [13].
In one aspect, a method is provided for treating a surface of a biomedical device. The method comprises depositing a polyanionic microgel onto the surface of the biomedical device, and loading the deposited polyanionic microgel with a peptoid.
In another aspect, a biomedical device is provided which comprises a surface selected from the group consisting of metal surfaces, ceramic surfaces and polymeric surfaces. A polyanionic microgel is disposed on said surface, and a peptoid is disposed in said polyanionic microgel.
In a further aspect, a self-defensive surface is provided which comprises a substrate; a polyanionic microgel disposed on said substrate; and a peptoid disposed in said polyanionic microgel.
In another aspect, a biomedical implant is provided. The biomedical implant comprises a substrate; and a release surface disposed on said substrate, wherein said release surface releases a peptoid in response to the presence of a pathogen; wherein said release surface includes a 3D cross-linked colloidal structure.
In a further aspect, a method is provided for treating a surface of a biomedical implant to render it resistant to bacterial cultivation. The method comprises (a) priming a surface of the implant with poly(allylamine hydrochloride), thereby obtaining a primed surface; (b) applying a sub-monolayer of poly(acrylic acid) (PAA) microgels to the primed surface, thereby obtaining a 3D cross-linked colloidal structure; and (c) loading the colloidal structure with a peptoid by forming a complex between the PAA microgels and the peptoid.
Self-defensive surfaces [14] represent an emerging strategy that may be able to inhibit bacterial colonization due to OR contamination. The term “self-defensive” was introduced by Boulmedais et al. [15] for the case of bacteria-triggered antimicrobial release. The mechanism at play with self-defensive surfaces is fundamentally different from conventional methods for inhibiting bacterial colonization through elutive drug delivery. In elutive techniques, antimicrobial compositions are continuously released from a surface whether they are needed or not. In contrast, a self-defensive mechanism releases the antimicrobial composition only when and where it is needed, and only in localized, very small amounts. Significantly, in the absence of a microbial challenge, self-defensive surfaces do not release antimicrobial compositions at all.
Self-defensive surfaces have been developed by Sukhishvili et al., [16] who designed polyelectrolyte thin-film coatings to release certain antimicrobial compositions in response to local pH changes. Cado, Boulmedais et al. [15a] exploited microbial enzyme secretion, and designed coatings that incorporate substrates for those particular enzymes, so that local coating degradation would release an antimicrobial composition. Both approaches are reliant on bacterial metabolism in order to become active. However, prior to implantation, a contaminated device surface has no source of nutrients, and thus, the occurrence of significant bacterial metabolism is unlikely. A self-defensive approach that relies on metabolism may thus not be appropriate for addressing OR contamination.
An alternate self-defensive mechanism, referred to as “contact transfer” [14c], has been developed and may be better suited to the problem of OR contamination. It has been shown [17] that contact transfer responds to the presence of challenging bacteria even in the absence of bacterial metabolism. In so doing, contact transfer drives the responsive release of certain antimicrobial compositions from a polyelectrolyte microgel coating, and can transfer those antimicrobial compositions to the challenging bacterium. This transfer has been attributed to the presence of a high concentration of negative charge on the bacterial envelope and the hydrophobicity of that envelop, which results in the complexation strength between the bacterium and the antimicrobial composition being greater than that between the microgel and the antimicrobial composition. Bacterial proximity to the loaded microgel then creates a steep chemical-potential gradient to drive antimicrobial de-complexation from the microgel and its transfer to the bacterium.
It has now been found that antimicrobial peptoids may be utilized in certain contact transfer surfaces to create effective self-defensive synthetic surfaces. Such surfaces may be fabricated, for example, by electrostatically depositing microgels of poly(acrylic acid) (PAA) to form a sub-monolayer coating on, for example, glass cover slips or titanium rods. In a subsequent self-assembly step, the sub-monolayer coating may be loaded with a suitable cationic antimicrobial peptoid. These self-defensive surfaces may significantly inhibit bacterial colonization, even in the absence of nutrients for metabolism. As described herein, the efficacy of these surfaces has been verified by testing in an in vitro model of OR contamination.
Various peptoids may be utilized in the self-defensive surfaces disclosed herein, although the use of the peptoid referred to as TM1 (also known as peptoid 1) is preferred.[18] Peptoids are analogs of peptides in which the side-groups are located on the amide nitrogen rather than on the α-carbon. In general, they are not proteolyzed,[19] thus improving their biostability and reducing immunogenicity. [19a] TM1 is a helical and amphipathic 12-mer peptoid consisting of a trimer (NLys-Nspe-Nspe) repeated 4 times and terminated with a secondary amine group (see
In situ optical microscopy may be utilized to track the peptoid loading process by imaging the microgel deswelling during complexation with the peptoid. Such imaging can also assess the ability of the peptoid to remain sequestered within the PAA microgels when exposed to physiologically relevant media. A digitally controlled aerosolizing system may be utilized to spray well-defined quantities of bacteria such as staphylococci (either S. aureus or S. epidermidis) onto Ti rods. Subsequent in vitro assays of bacterial viability and osteoblast response show that the peptoid-loaded microgel-modified surfaces inhibit bacterial colonization very effectively while still preserving cytocompatability comparable to unmodified controls.
Poly (acrylic acid) (PAA) microgels were synthesized by thermally initiated membrane emulsification. A precursor solution was mixed using 1.0 ml acrylic acid (Sigma), 0.47 g sodium hydroxide (NaOH, Sigma), 4 ml deionized (DI) water (Millipore type 1), 100 mg ammonium phosphate sulfate (APS, Sigma) and 100 μl Poly(ethylene glycol) diacrylate (PEGDA, Mn=575 Da). This aqueous solution was forced via N2 pressure (40 kPa) through a ceramic membrane (1.5 μm pore size (Shirasu Porous Glass (SPG)) into a stirred (400 rpm) oil phase consisting of 2.56 g Span 80 and 160 ml paraffin oil. The resulting emulsion was then deoxygenated for 30 min by N2 bubbling followed by heating to 70° C. and held there for 4 h under continuous stirring (500 rpm). After cooling to room temperature, the paraffin oil was removed by centrifugation and re-suspension first in cyclohexane two times, then in ethanol ten times, and finally in DI water ten times. The resulting PAA microgels were suspended in sterilized DI water and stored at 4° C.
For in situ studies of microgel/antimicrobial interactions, a polydimethylsiloxane (PDMS) gasket was used to define 12 individual reaction chambers above a glass microscope slide. 6 mm diameter holes were punched from a cast sheet of cured PDMS (˜4 mm thickness). The punched gasket was then pressed onto a pre-cleaned (3 min oxygen plasma) glass slide and annealed at 70° C. for 20 min. The volume of each glass-bottomed chamber was ˜100 μl. The glass surface within each chamber was primed with positively charged poly(allylamine hydrochloride) (PAH, Sigma, Mw=17.5 kDa) using an aqueous solution of 0.2 mg PAH/ml for 1 h followed by gentle washing using DI water and then drying by flowing N2 gas. A sub-monolayer of PAA microgels was then electrostatically deposited onto the primed glass surface by filling each chamber with a colloidal aqueous microgel suspension at room temperature and left to soak for 30 min. These chambers were then rinsed with DI water and filled with 0.01 M phosphate buffer (pH 7.4, 4 mM NaH2PO4 and 6 mM Na2HPO4), [Na+]=0.016 M) prior to subsequent antimicrobial loading.
TM1 (Mw=1819 Da) was synthesized as previously described.[21] It was loaded by complexation within the microgels by replacing the buffer within a particular reaction chamber with a solution of 1 mg TM1/ml of 0.01 M phosphate buffer. The loading process was followed by in situ time-resolved optical microscopy using an inverted microscope (Nikon, Eclipse Ti-E) equipped with a 14 bit CCD Camera (pco.pixelfly) and a CFI S Fluor ELWD 20× objective lens (NA=0.45, WD=8.2-6.9 mm). During the first 10 minutes of loading images were collected at one-minute intervals and then at five-minute intervals thereafter. After data acquisition, the time-resolved diameters of at least five different microgels were measured using ImageJ.[22] Each diameter was normalized using the initial diameter measured in TM1-free 0.01 M phosphate buffer.
The sequestration of the TM1 complexed within the microgels was again followed by in situ imaging. After loading, the TM1 solution was removed by gentle washing using DI water (3 times). Each reaction chamber was then filled with either autoclaved 0.01 M phosphate buffer or autoclaved PBS (phosphate buffered saline, pH 7.4, ionic strength=0.138 M). The buffer was replaced with fresh buffer daily for 28 days. Optical micrographs of the same set of microgels were collected daily, and digital image analysis was used to measure the microgel diameter as a function of soaking time from at least five different microgels.
The amount of TM1 loaded into the microgels was quantified by UV absorption. Two aliquots of 10 μl of aqueous PAA microgel suspensions were prepared. One was allowed to fully dehydrate to measure the dry microgel weight. The other was added to 1 ml 0.01 M phosphate buffer containing 0.25 mg of TM1. After 60 min the loaded microgels were pelletized by centrifugation, and the UV adsorption (230 nm) of the supernatant was measured using a Synergy HT BioTek Spectrometer. The amount of loaded TM1 was then quantified by comparison with a calibration curve created using solutions of known TM1 concentrations in 0.01M phosphate buffer. The microgel zeta potential was measured using a Malvern ZETA SIZER Nano series from aliqouts of microgels suspended in pure 0.01 M phosphate buffer and before/after TM1 loading.
Commercially pure, surgical grade 1, titanium rods (1 cm length) were cut from wire (Temco RW0469) with an average diameter of 1.29 mm. The rods were rinsed twice with 70% ethanol, sonicated in ethanol (15 min), washed and sonicated in autoclaved DI water, and finally dried using flowing N2 gas. The dried rods were then exposed to an oxygen plasma for 10 min. Further modifications were made to create four surfaces: (Ti-A) plasma treated but otherwise unmodified Ti; (Ti—B) PAH-primed Ti; (Ti—C) microgel-modified (unloaded) Ti; and (Ti-D) TM1-loaded microgel-modified Ti. The PAH priming, electrostatic PAA microgel deposition, and the TM1 loading all followed the procedures described above for the glass substrates. The surfaces of each condition were imaged by scanning electron microcopy (SEM; Zeiss Auriga). TM1 loading for condition D was confirmed by exposing the modified pins to an aqueous solution of fluorescein isothiocyanate (FITC, Sigma) for 30 min in sterile DI water followed by repeated washing with DI water. The FITC-stained Ti-D rods were then imaged while hydrated using the Nikon Eclipse Ti-E inverted microscope with a C3 confocal attachment.
A digitally controlled aerosolizing system was used to contaminate the surfaces of rods in each of the four surface conditions with well-controlled quantities of bacteria. Details of the system configuration, operation, and properties have been described elsewhere.[23] The system is able to reproducibly spray well-defined bursts of aerosolized bacteria suspended in nutrient-free buffer onto test surfaces. Depending on the spray conditions, the density of sprayed bacteria is on the order of 102-103/cm2 and can thus mimic the contamination conditions of a surgical theater. The experiments here used methicillin-sensitive Staphylococcus aureus (MSSA, ATCC 29213) and Staphylococcus epidermidis (ATCC 35984). For each set of experiments, a single bacterial colony was inoculated into 10 ml of Tryptic Soy Broth (TSB) and grown to stationary phase at 37° C. under gentle shaking for 18 h. The bacteria were pelleted by centrifugation and resuspended in sterile PBS twice to remove the TSB. After final pelleting, the bacteria were suspended in sterile PBS to an optical density (OD600) of 0.0005 A for MSSA and 0.0004 A for S. epidermidis. These densities have been shown to produce contamination levels of about 102 CFU/cm2.[23] The solutions were kept at room temperature and used within 3 hours of preparation. It has been shown [23] that over 90% of the bacteria remain culturable under these preparation conditions.
Two sets of three Ti rods from each of the four surface-modification conditions were sprayed with 100 ms bursts of bacterial aerosol. One set was sprayed with MSSA, and the other was sprayed with S. epidermidis. The three rods from each set were sprayed simultaneously. After spraying, the samples were left to dry in air within a laminar-flow biosafety hood for 30 min. The extent of contamination was then quantified in terms of the numbers of: (i) loosely bound culturable bacteria and (ii) strongly adhered culturable bacteria. Loosely bound bacteria were recovered by immersing each rod in 1.25 ml of sterile PBS followed by gentle vortexing. Three 250 μl aliquots of the buffer were then spread on separate tryptic soy agar (TSA) plates. Strongly adhered bacteria were assayed by gently washing each vortexed Ti rod using sterile PBS and then rolling each rod using flame-sterilized forceps over the surface of a TSA plate. Both sets of TSA plates were then cultured at 37° C. for 18 h. The number of colony-forming units (CFUs) was counted manually. The statistical significance of differences in bacterial colonization of the different surface-modification treatments was determined using a two-sided student t-test assuming unequal variances for the treated and control rods.
Human fetal osteoblasts (hFOB, ATCC, VA) were cultured in a 1:1 mixture of Dulbecco's Modified Eagle's Medium (DMEM) and Ham's F12 Medium, containing 10% fetal bovine serum (FBS, Atlantic Biologicals) and 2.5 mM L-glutamine (without phenol red). Prior to use, the culture was maintained at 34° C. with 5% CO2 and 95% humidity until 70-80% confluence. To achieve a homogeneous cell distribution on individual substrates (Ti rods or glass cover slips) and equally among multiple substrates, hFOB cells were seeded and cultured following a specific protocol. Briefly, the Ti rods and coverslips sampling each of the four surface-modification conditions (A-D) were sterilized using 70% ethanol and then washed for 2 min with sterile DI water (3×). All samples were again sterilized under UV irradiation for 20 min before cell seeding. Each rod was first seeded with 100 μl of hFOB cell suspension, and the cell-seeded rod was then incubated for 15 min (34° C., 5% CO2, and 95% humidity) to promote cell attachment. The seeded rod was then flipped over and exposed again to 100 μl of hFOB cell suspension to achieve a final seeding density of 2×104 cells/cm2. Coverslips were seeded with 200 μl of hFOB cell suspension to a seeding density of 104 cells/cm2. After 30 min of total incubation, all cell-seeded samples were transferred into a new 6-well plate for further cultivation. The culture medium was refreshed every 2 days.
The adhesion (1 day) and proliferation (4 day and 7 day) of hFOB cells on the rods were assessed using an MTS assay. Briefly, after 1, 4, and 7 days, the culture rods (n=3 for each condition) were harvested, placed in a 12-well plate, and gently rinsed with cold PBS twice. Then, 1.4 mL of a mixture containing 400 μL MTS reagent (the assay kit was purchased from Promega, Madison, WI) and 1 mL culture medium was added each well. The plate was incubated for 2 h at 37° C. in 5% CO2. 200 μL of supernatant from each sample was transferred to a 96-well plate. The absorbance at 490 nm was recorded using the BioTek Synergy microplate reader (BioTek Instruments, Inc., Vermont, USA). Statistically significant differences were assessed using a one-way t-test for the case where the Ti-A surface was used as the control and using an ANOVA test by Tukey's mean comparison (p<0.05) for the case where the Ti-A, B and C surfaces were used as controls.
The cell morphologies on Ti wires and coverslips were characterized after cultivation of 1, 4, and 7 days. The specimens were fixed with 4% (w/v) paraformaldehyde for 15 min and washed for 2 min with PBS (3 times). Then, the samples were rinsed 3 times with 0.5% Triton X-100 and blocked with 3% (w/v) BSA to permeabilize the cell membranes. The cell nuclei were stained with 4′, 6-diamidino-2-phenylindole (DAPI, Sigma, USA), and the cell skeletal filament actin (F-actin) was stained with Alexa Fluor 488-conjugated phalloidin. After staining, the samples were immersed in PBS and imaged while hydrated using a Nikon eclipse 80i epifluorescence microscope (Japan) and Zeiss LSM 880 confocal laser scanning microscope (Germany).
When hydrated in 0.01 M phosphate buffer at pH 7.4 and [Na+]=0.016 M, the as-synthesized PAA microgels have an average diameter of 6±2 μm (
Since device colonization is possible not only prior to implantation when the device is dry but also prior to surgical-site closure, an important question centers on whether the TM1 remains stably complexed within the PAA microgels when the device is in contact with physiological fluids. At a constant pH of 7.4, when the ionic strength is increased from [Na+]=0.016 M in the 0.01 M phosphate loading buffer to [Na+]=0.14 M typical of physiological conditions, the additional salt can shield the electrostatic pairing between amine groups on the TM1 and the acid groups within the PAA microgels. Such shielding, for example, leads to the burst release of colistin from PAA microgels.[17, 20] The fact that
Having developed a means to create TM1-loaded PAA-microgel modified surfaces by directed self-assembly, the contamination resistance of Ti rods subjected to each of four different surface treatments was investigated. These are illustrated schematically in the top panel of
To mimic intraoperative contamination by airborne bacteria, three Ti rods were sprayed, sampling each of the four surface conditions with aerosols of 0.0067 M phosphate buffer (1×PBS with 0.0067 M PO4, pH 7.0 to 7.2) containing either MSSA or S. epidermidis. After spraying, the rods were left to dry and then assayed to determine the number of culturable bacteria either loosely bound or strongly adhered to the Ti surface. These two cases are differentiated mechanically by vortexing each rod in 0.0067 M phosphate buffer. Loosely bound bacteria were recovered in the buffer, and strongly adhered bacteria were recovered after vortexing by rolling the contaminated rod on an agar plate.
Importantly, the number of culturable bacteria found on the TM1-loaded microgel-modified Ti-D samples is very small. In three trials, an average of only 2 CFUs of MSSA and only 1 CFU of S. epidermidis were found. These few CFUs were all recovered from the buffer after vortexing the contaminated rod and thus correspond to loosely bound colonies. No CFUs were found from bacteria strongly adhered to the Ti-D rods.
Since the three control surfaces all become significantly colonized, it is assumed that the Ti-D rods gare similarly exposed to culturable bacteria as a result of the spray-contamination process. The fact that only a very small number of colonies is subsequently found on the Ti-D rods indicates that the majority of these bacteria were killed by interaction with the surface. Neither the surface nor the medium (0.067 M phosphate buffer) in which the bacteria were suspended during the aerosolizing process provide nutrients, so it is expected that the bacteria will undergo little or no metabolism on the surface. Hence, triggering TM1 release from the loaded microgels via local pH reduction [25] is unlikely. Instead, these findings support the concept of contact transfer where the local chemical potential of the TM1 is lower in the bacterial envelope than it is within the microgels, [14c, 17] so that, when a bacterium comes within close proximity, some TM1 de-complexes from the microgel and re-complexes within the staphylococcal cell envelope. Such complexation interactions with bacteria have been extensively studied for a number of host-defense peptides and are attributed to the rich high concentration of anionic, hydrophobic, and aromatic moieties in the outer bacterial membranes.[26]
The fact that microgels, even when loaded, remain hydrophilic promotes opportunities for microgel-bacterium contact. If an aerosolized droplet impinges on a dry microgel-modified surface, water from that droplet will locally hydrate one or more microgels. As the overall droplet evaporates, the microgel(s) will be the most hydrophilic element of the surface and will be last to fully dry. Without wishing to be bound by theory, it is believed that surfaces forces associated with the edge of the drying droplet will draw any bacteria it may contain to a microgel, and when droplet evaporation is complete those bacteria will be concentrated on or near a microgel.
It is also notable that the in vitro model of OR contamination does not overwhelm the surface with bacteria as can be the case with in vivo infection models where a concentrated inoculum (e.g. 106-108 CFU/mL) is typically required in order to provoke infection in control animals such as rats and mice. The OR contamination model sprays low numbers of bacteria onto a rod surface. The probability of any particular microgel interacting with a bacterium is low. For that (much larger) subset of microgels that are unchallenged, the antimicrobial remains sequestered within these microgels (see the SEM image of the ToD rod in
Cell morphology (imaging) and metabolic activity (MTS) were utilized to evaluate the short-term cytocompatibility of the various surfaces. In the case of microgel-modified surfaces, only a fraction of the surface is covered by microgels. The underlying PAH-primed Ti surface is exposed between adjacent microgels and, because the PAH is cationic, it is expected that this exposed surface becomes covered by negatively charged serum proteins (e.g. albumin, fibronectin, etc.) when exposed to serum-containing medium. It has been previously shown that such microgel-modified surfaces remain highly compatible with both osteoblasts and macrophages.[17, 27] It is notable that the unmodified Ti surface has an intrinsic roughness over lateral length scales on the order of 2-10 μm, and that this roughness is further perturbed by the addition of TM1-loaded microgels with characteristic dimensions on the order of several microns. Such surface topography has long been known to affect and, in many cases, enhance cell adhesion, spreading, and proliferation.[28]
Glass substrates were utilized to avoid many of the imaging complications associated with the curved and rough surface of the Ti rods. The top images in
Exposure to the OR atmosphere between the time when a medical device is removed from its sterile package and the time when the wound site is fully closed can lead to device contamination by bacteria sedimenting from the atmosphere. That process has been modeled using an aerosolizing system able to spray small quantities of bacteria onto titanium rods. Even though relatively low numbers of bacteria contaminate each rod, when exposed to culture medium, bacteria sprayed onto unmodified Ti rods develop into proliferating colonies. The experiments described herein show that Ti surfaces modified with polyanionic microgels loaded by complexation with a cationic antimicrobial peptoid (preferably TM1) are able to almost entirely inhibit bacterial colony formation. Nutrients are unavailable to enable bacterial metabolism during the contamination process, indicating that the bacteria trigger local TM1 release by contact transfer rather than by local changes in pH. Because of the low numbers of contaminating bacteria, most of the loaded microgels are unchallenged during contamination, so their TM1 payload remains sequestered. However, neither the loaded peptoid nor the additional topography introduced by the microgels diminish the cytocompatibility of the modified surfaces as assayed by in vitro experiments with human fetal osteoblasts. This surface-modification strategy thus suggests a promising approach with which to inhibit the intra-operative bacterial colonization of exposed biomedical devices due to contamination in the operating theater.
While frequent reference has been made herein to TM1 (H-(NLys-Nspe-Nspe)4-NH2), various peptoids may be utilized in the devices, methodologies and compositions disclosed herein. Suitable peptoids may include, for example, the peptoids described in U.S. Pat. No. 8,445,632 (Barron et al.), entitled “Selective Poly-N-Substituted Glycine Antibiotics”, which is incorporated herein by reference in its entirety, and the peptoids having the designations and structures depicted in TABLES 1-2 below.
Various cyclic peptoids may also be utilized in the devices, methodologies and compositions disclosed herein including, but not limited to, the peptoids disclosed in U.S. Pat. No. 9,938,321 (Kirshenbaum et al.), U.S. Pat. No. 9,315,548 (Kirshenbaum et al.) and U.S. Pat. No. 8,828,413 (Kirshenbaum et al.), all of which are incorporated herein by reference in their entirety.
In some embodiments, the peptoids disclosed herein are a poly-N-substituted glycine compound of a formula
The peptoids in the compositions described herein may be alkylated, and preferably have terminal alkylation. Here, alkylation (and especially terminal alkylation) with a C10 or C13 tail is especially preferred. It has been found that such terminal alkylation may enhance the efficacy of the peptoid in some applications, and in some cases, may cause a peptoid which otherwise has low antibacterial activity to have significant antibacterial activity. Thus, in the foregoing formula, R may be selected from about C4 to about C20 linear, branched and cyclic alkyl moieties. R preferably contains at least 8 carbon atoms, more preferably contains at least 10 carbon atoms, and most preferably contains at least 12 carbon atoms. Embodiments where R is a decyl or tridecyl moiety are especially preferred.
It will further be appreciated that various salts or precursors of the foregoing peptoids may be utilized in the compositions and methodologies disclosed herein. As an example of the latter, these compositions and methodologies may contain one or more precursors of the general formula A-P, where P is one of the foregoing peptoids (or a salt thereof) and A is a moiety which acts as a leaving group to generate P when the precursor undergoes an in vivo reaction (such as, for example, proteolytic degradation).
Various halogenated peptoids and halogenated oligomers of N-substituted glycines (and salts thereof) may also be utilized in the compositions and methodologies disclosed herein. These include, without limitation, various halogenated analogs of the foregoing peptoids and oligomers of N-substituted glycines. Examples of such halogenated peptoids are described, for example, in WO2020223581A1 (Molchanova et al.), entitled “Halogenated Antimicrobial Peptoids”, which is incorporated herein by reference in its entirety. These halogenated compositions may be halogenated in various ways. For example, these compounds may include any number of halogen substitutions with the same or different halogens. In particular, these compounds may include one or more fluoro-, chloro-, bromo- or iodo-substitutions, and may include substitution with two or more distinct halogens. However, the use of one or two bromo- or chloro-substitutions is preferred in many applications. Moreover, while the peptoids described herein may be halogenated at various locations, para halogenation on the peptoids containing aryl rings is especially preferred in many applications, although ortho- and meta-substitution, or even perhalogentation, may be useful in some applications.
Any of the peptoids disclosed in [Nam, H. Y., Choi, J., Kumar, S. D., Nielsen, J. E., Kyeong, M., Wang, S., Kang, D., Lee, Y., Lee, J., Yoon, M.-H., Hong, S., Lund, R., Jenssen, H., Shin, S. Y., Seo, J., 2020. Helicity Modulation Improves the Selectivity of Antimicrobial Peptoids. ACS Infectious Diseases 6, 2732-2744. doi:10.1021/acsinfecdis.0c00356], which is included herein by reference in its entirety, may also be utilized in the compositions and methodologies disclosed herein. These specifically include the peptoids depicted in TABLE 2 below.
Various biomedical devices may be treated in accordance with the teachings herein. These may include, for example, hip/knee prostheses, heart valves, pacemakers, cochlear implants, shunts, surgical mesh, sutures, and tissue-engineering constructs. These biomedical devices may be of as type that are designed to be implanted in the body of a subject, which may be a human or animal subject.
Preferred embodiments of the methodologies disclosed herein involve treating a surface of a biomedical device by depositing a polyanionic microgel (or in some embodiments, a nanogel) onto the surface of the biomedical device and loading the deposited polyanionic microgel with a peptoid. In some embodiments, the surface of the biomedical device may be primed with a polycation layer prior to depositing the polyanionic microgel. The polyanionic microgel may be synthesized by a process which includes membrane emulsification and UV photopolymerization and is preferably deposited as a sub-monolayer coating on the surface of the biomedical device. Preferably, the polyanionic microgel comprises polyacrylic acid.
The surfaces of biomedical devices treated in accordance with the methodologies disclosed herein may comprise various materials. These include, without limitation, materials selected from the group consisting of metals, metal oxides, and metal alloys such as, for example, alumina, zirconia, and zirconia toughened alumina (ZTA); refractory metals, such as Zr, Ta, V, Nb, W, Mo and alloys thereof; metal alloys, such as, for example, Ti-6Al-4V alloys, Ti-5Al-2.5Fe alloys, Ti-6Al-7Nb alloys and β-Ti alloys (including Ti-12Mo-6Zr-2Fe (TMZF)); titanium alloys, such as, for example, Co—Cr—Mo alloys, stainless steel alloys, and refractory metal alloys; oxidized zirconium (oxinium); alumina, zirconia, and zirconia toughened alumina (ZTA); magnesia partially stabilized zirconia (MgPSZ); and yttria stabilizing oxide (Y-TZP). The surfaces of biomedical devices treated in accordance with the methodologies disclosed herein may also comprise various organic polymers such as, for example, polyethylene (including ultra-high-molecular-weight polyethylene (UHMWPE) and highly cross-linked polyethylene (HXLPE)), polytetrafluoroethylene (PTFE) or polyetheretherketone (PEEK).
The self-defensive surfaces disclosed herein may be adapted to release one or more peptoids in the presence of various pathogens, including various bacteria, fungi and viruses. Specific examples of such pathogens may include Staphylococcus aureus, Pseudononas aeruginosa, Escherichia coli, Klebsiella, Proteus, Enterobacter, Clostridium difficile, and Salmonela, Streptoccoci; Candida albicans, Aspergillis spp., Nocardia, Pneumocystis carinii, Cryptococcus neoformans, and Cryptosporidium; respiratory syncytial virus, cytomegalovirus, human immunodeficiency virus (HIV), Ebola, rotavirus, enteroviruses, Influenza A (including subtypes H2N2 and H3N3), hepatitis and herpes viruses.
Various peptoid release surfaces may be created in accordance with the teachings herein. These release surfaces preferably comprise microgels, but in some embodiments may also comprise nanogels. The release surface preferably comprises a zwitterionic polymer, such as a polybetaine. Such a polymer may contain cationic groups such as, for example, quaternary ammonium groups; and anionic groups such as, for example, carboxylate, sulfonate, phosphate, phosphinate, and phosphonate groups.
In some embodiments, methods are provided for treating a surface of a biomedical implant to render it resistant to bacterial cultivation. In such embodiments, the method may comprise (a) priming a surface of the implant with poly(allylamine hydrochloride), thereby obtaining a primed surface; (b) applying a sub-monolayer of poly(acrylic acid) (PAA) microgels to the primed surface, thereby obtaining a 3D cross-linked colloidal structure; and (c) loading the colloidal structure with a peptoid by forming a complex between the PAA microgels and the peptoid. The poly(allylamine hydrochloride) is preferably positively charged, and treating a surface of the implant with poly(allylamine hydrochloride) may include (a) applying an aqueous solution of poly(allylamine hydrochloride) to the surface, thereby forming a treated surface; (b) rinsing the treated surface, thereby obtaining a rinsed surface; and (c) drying the rinsed surface. Applying a sub-monolayer of PAA microgels to the primed surface may include electrostatically depositing PAA onto the primed surface, which may include exposing the primed surface to a colloidal aqueous microgel suspension. Similarly, loading the PAA microgels with a peptoid may include exposing the PAA microgels to a buffered solution of the peptoid, in which case the use of a phosphate buffer is especially preferred. The loading process preferably includes forming a complex between the PAA microgels and the peptoid.
PAA microgels of various dimensions may be utilized in the constructs and methodologies described herein. Preferably, the PAA is have an average diameter within the range of 6±2 μm prior to loading the PAA microgels with a peptoid. Loading the PAA microgels with a peptoid preferably results in a ratio Raa/pep of acrylic acid groups to peptoid within the range of 1 to 50, more preferably 2 to 20, even more preferably 3 to 10, and most preferably 4 to 5. The microgel zeta potential preferably increases from a value of −33.5±4.5 mV in the unloaded state to 3.6±1.2 mV in the loaded state.
The above description of the present invention is illustrative, and is not intended to be limiting. It will thus be appreciated that various additions, substitutions and modifications may be made to the above described embodiments without departing from the scope of the present invention. Accordingly, the scope of the present invention should be construed in reference to the appended claims.
The present application claims priority to U.S. provisional application 63/214,782 (Libera et al.), entitled “SUBSTRATES MODIFIED WITH PEPTOID-LOADED MICROGELS FOR RESISTANCE TO BACTERIAL COLONIZATION”, which was filed on Jun. 25, 2021, and which is incorporated herein by reference in its entirety.
This invention was made with government support under grant #W911NF2010277 awarded by the Army Research Office; grant #DMR-1608406 awarded by the National Science Foundation; grant #1DP1 OD029517-01 awarded by the U.S. Public Health Services; and grant No. DE-ACO2-05CH11231 awarded by the U.S. Department of Energy.
| Filing Document | Filing Date | Country | Kind |
|---|---|---|---|
| PCT/US2022/035189 | 6/27/2022 | WO |
| Number | Date | Country | |
|---|---|---|---|
| 63214782 | Jun 2021 | US |
