Local retention of antibodies in vivo with an injectable film embedded with a fluorogen-activating protein
Graphical abstract
Introduction
Monoclonal antibodies are an expanding class of biotherapeutics [1]. While antibodies are typically formulated for intravenous or intramuscular administration, localized delivery can concentrate the biotherapeutics in diseased tissues, thereby increasing bioavailability to an extent that cannot be achieved through systemic routes [2]. Subcutaneous or intramuscular delivery can target regional lymph nodes in order to modulate local immune milieu [3]. Outcomes of inflammation are governed in part by the phenotypes of granulocytes, myeloid cells, and lymphocytes trafficking between injured tissues and draining lymph nodes. In situ-forming hydrogels are used to create depots of antibody drugs in vivo. Local administration of immune modulators can enhance efficacy and reduce adverse events. In-situ forming hydrogels or coacervates formed by cross-linked polymer chains that remain physically intact with high water intake are particularly attractive because of their ability to adapt to tissue space [2]. Drug loading can be accomplished by introducing the biologic during polymer cross-linking. The pore sizes (5–100 nm) and the resulting high permeability of such depots [4], however, preclude the ability to concentrate antibodies at injection sites and accumulating in draining lymph nodes. Rapid diffusion of drugs out of the network (“burst release”) can be limited by increasing cross-linking density [5] or through bioaffinity retardation [6]. A less explored challenge is that the location of the depot cannot be ascertained in vivo, making it difficult to determine the rate-limiting mechanism of drug localization.
The EAK16 series of β-sheet fibrillizing peptides (βFP) are used to formulate depots of small molecule drugs, proteins, and cells [7], [8], [9]. The βFP EAK16-II, comprising the sequence AEAEAKAKAEAEAKAK (with A as alanine, E as glutamic acid, and K as lysine), self-assembles into cross-linking fibrils in high ionic strength aqueous environment [7]. When administered through conventional needled syringes, EAK16-II rapidly forms stable coacervates in vivo, adapting and adhering to the local tissue contours [10]. Due to the high water content (~ 99.5% w/v) [7], IgG can be loaded into the fibrillar matrices with high efficiency. The amphiphilic EAK16-II, carrying no net charge at pH 7.4, presumably interacts with IgG via non-specific interactions, resulting in a large fraction of the protein molecules remaining unbound [11]. In hydrogels made with the βFP sequence RADA, drug release kinetics is governed by the molecular size of the encapsulated protein; IgG, the largest molecule tested, reaches maximum release by 24 h [12]. Koutsopoulos et al. subsequently demonstrated that burst release can be suppressed by encapsulating IgG in double-layer nanofibers using two amphiphilic building blocks, RADA and KLDL [13]. These studies demonstrate the versatility of βFP in controlling the release of protein drugs.
Biomolecules provide affinity mechanisms for tuning protein release from hydrogel matrices. The Sakiyama–Elbert group has engineered heparin-embedded hydrogels as protein delivery systems. As highly sulfated glycosaminoglycans, heparins bind growth factors through ionic reversible interactions. These systems mimic reservoirs of endogenous growth factors in extracellular matrix [14], [15], [16], [17], [18], [19] in which secreted proteins are protected from denaturation and aggregation. Hydrogels functionalized with aptamers have been used to control protein release [20], [21], [22], [23]. Diverse sequences of aptamers can be generated to confer different ligand specificities and affinities. The Shoichet group has used Src homology domain 3 (SH3) to tune the release of growth factors fused with SH3 ligands [24], [25], [26], [27]. In these designs, rapid diffusion is suppressed; instead, drug release is a function of the affinity of the binding interactions (KD), the relative ratio of binding sites (ligands) to proteins, and the dissociation rate constant (koff) of the protein-binding site complexes [26]. Thus affinity-based hydrogels afford the ability to tune release kinetics by adjusting classical biochemical parameters.
We have used the coassembly method to engineer biochemical functionalities into EAK16-II fibrillar matrices [10], [28], [29], [30], [31]. Using the analog EAKIIH6, His-tags are incorporated into EAK16-II fibrils, driven by the βFP domain for β-sheets. The two peptides co-assemble efficiently, with greater than 90% of EAKIIH6 incorporated into the composite [10] generated from excess EAK16-II. By embedding recombinant protein A/G (pAG) via anti-His-tag antibodies, IgG can be loaded and release over time [29], [30], [31]. Because Protein A and Protein G bind to IgG with relative high affinities (with apparent binding constants reported at 2.9 × 107 and 1.13 × 108 M− 1 [32], respectively), pAG provides a stable affinity mechanism, allowing diffusive release of the IgG in vivo driven by the concentration gradient between the depot and the open system of surrounding tissues and lymphatics. Staphococcal Protein A (SpA) has been developed into a systemic therapeutic agent (PRTX-100) recently approved for use in humans [33]. The system renders rapid formation of depots and retention of IgG molecules in subcutaneous injection sites for up to ten days [29]. The EAK16-II/EAKIIH6 system, when adapted with pAG and anti-His-tag antibodies, was designed as antibody-functionalized scaffolds for cytokines and cells [29], [30]. The ability to monitor IgG-binding sites in vivo will aid efforts in correlating in vitro and in vivo observations.
Herein we described an in situ-forming film on which the binding sites for IgG can be monitored using an activation-dependent fluorescent module. The films are generated by intermixing EAK16-II with its analog dL5_EAK, a polypeptide consisting of a FAP with two light chains, called dL5, fused to three repeats of AEAEAKAK via a glycine–serine spacer, (GGGGS)3 [28]. In high ionic strength environment, EAK16-II and dL5_EAK undergo solution-film transition and coassembly, apparently simultaneously. The resultant materials are illuminated in the presence of malachite green (MG), the fluorogen activated by dL5 [34]. Relative to free MG, the fluorescence of dL5-bound MG is enhanced by up to 20,000 folds, due to constraining of the rotatable bonds in MG by two L5 proteins. Crystallographic studies have shown that MG mediates dimerization of two L5 proteins to form a fluorescent ternary complex [35]. We hypothesized that the system of MG-conjugated pAG (pAGMG), dL5_EAK and EAK16-II, when deployed together (called dL5 film), serves as a self-illuminating, injectable formulation for local deposition of IgG. The tight binding of dL5 to MG ensures a stable intermediate [35]. Release of IgG from pAGMG is driven by the diffusion of free IgG into tissues surrounding the injection site. The exceptionally low fluorescence background renders the fluorogenic module an efficient reporter of binding sites, independent of model IgG drugs [34]. We characterized and modeled the release kinetics in vitro, and found that the dL5 film retained IgG at the injection site and draining lymph node in vivo.
Section snippets
EAK16-II and expression of dL5_EAK
EAK16-II was custom synthesized by American Peptide Company (Sunnyvale, CA) at greater than 95% purity with the N-terminus acetylated and C-terminus amidated. The sequence was confirmed by mass spectrometry by the manufacturer. The peptide, received lyophilized, was reconstituted in sterile deionized water (18.2 MΩ at 25 °C) as needed and stored at − 80 °C until use. Construction of plasmids encoding dL5_EAK was described previously [28]. Expression and purification of dL5_EAK was carried out by
Results
We developed a bioaffinity-driven system in which the release of IgG from a fibrillar matrix was retarded in order to concentrate therapeutic antibodies locally at injection sites and in regional lymph nodes. By concentrating antibodies in lymph nodes to which inflamed tissues drained, dosage of the drug injected can be reduced. Techniques of parental injections targeting selected lymph nodes have been reported [41]. The current design hinges on engineering a fluorogenic module into the network
Discussion
This work was motivated by the need for a trackable injectable system by which antibodies could be concentrated in vivo. Potential applications include intratumoral antibody injection [29], and topical delivery of anti-cytokine antibodies [44]. The inflammatory mediator TNFα driving non-healing cutaneous wound is a case in point. In addition to releasing soluble TNFα, macrophages infiltrated into the wound tissues express transmembrane TNFα [45]. Together with neutrophils and adipocytes,
Conclusions
The results demonstrate biochemical and imaging functionalization of a βFP-based injectable platform. The depot platform can be used to deliver IgG locally of a variety of specificities. The delayed release demonstrated in vitro might be translated into in vivo prototypes through iterative optimization, guided by in situ detection of FAP-generated fluorescence. The design could be used to deliver antibodies for suppressing cutaneous inflammation and targeted immune modulation in draining lymph
Acknowledgments
This work was supported in part by a Health Research Formula award from the Pennsylvania Department of Health (W.S.M.), and National Institutes of Health grants R21AI113000 (W.S.M) and U54GM10329 (A.S.W.).
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2019, Acta BiomaterialiaCitation Excerpt :Inspired by systems of controlled release in polymeric gels modified with affinity ligands in heparin [15], divalent metal [16–18], and SH3 domain [19,20], we have engineered Fc-binding functions into EAK gels in formulating IgG for localized delivery [13,21,22]. The general strategy is to admix EAK with a second EAK-containing peptide fused with an affinity domain, such as His-tag [23,24] and dL5 [25,26]. The capture and release of IgG by these co-assemblies are governed in part by the equilibrium binding constant (KD) of IgG interacting with protein A/G, which serves as a linker between the gel matrix and the antibody.
Physical characterization and modeling of chitosan/peg blends for injectable scaffolds
2018, Carbohydrate PolymersCitation Excerpt :The flow of polymers is a complex process extensively studied and modelled considering physical structures from molecular arrangements of chains and how chains move or interact (Budtova & Navard, 2015). However, injectability has been only studied for pharmaceutical proteins or drugs (Liu et al., 2016; Yan et al., 2012) and even though scaffolds have found some increasing interest there is still few information about its most important feature that is injectability. It is considered that the ease or the force required to remove an injectable scaffold from the syringe must be determined for the finished product.
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Present address: Glycosensors & Diagnostics, JLABS, 3210 Merryfield Row, San Diego,
CA 92121, United States.