Protein Delivery
4.3.1 Strategies for Protein Release from Hydrogels
Proteins can be loaded into hydrogels through manifold mechanisms and strategies.
They can be physically entrapped in hydrogels or adsorbed into the matrix by spe-cific interactions, noncovalent or covalent binding via degradable linkers, and subse-quently released via diffusion, swelling, erosion, degradation or a specific trigger such as pH, temperature, etc. Also, a combination of multiple delivery systems can be used to achieve modular release of multiple proteins. Depending on the release mechanism, different release profiles can be obtained. Table 4.2 overviews the technologies further described in this chapter.
4.3.1.1 Physical Entrapment of Proteins into Hydrogels: General Principles and Release Mechanisms
Physical encapsulation of proteins in order to obtain a controlled release of the active is an easy and frequently applied technique. Its relative simplicity represents a clear advantage over more sophisticated encapsulation/release methods and the majority of hydrogel systems rely on this loading process. Loading is done by incubation of the preformed gel with the protein of interest or by adding the protein to the hydrogel forming monomers/prepolymers. In recent years, attention has shifted to in-situ gell-ing thermosensitive systems, which preferably are liquid before administration and jellify at the site of injection [26]. For these gels, loading with proteins can easily be established by dissolution of the biotherapeutic in the gel forming polymer solution before administration. Typically, the mechanisms that govern the drug release from these hydrogels are controlled by diffusion, swelling, erosion, external stimuli or combinations thereof [115]. The release mechanism depends on both the character-istics of the polymeric network and the protein. When the hydrogel pores are bigger than the hydrodynamic radius of the protein, diffusion is the driving force for release, with a diffusion rate depending on the protein size and the water content of the gel
Figure 4.6 Schematic of the concept of controlled protein delivery using injectable hydrogel.
110 Handbook of Polymers for Pharmaceutical Technologies
Table 4.2 Overview of the protein release mechanisms from hydrogels. Reproduced from [114].
Release mechanism from hydrogels Characteristics
Fickian diffusion
Mt/M∞ = k√t
Concentration gradient
Generally limited to relatively short time spans Protein size < mesh size
Geometry determines release rate
Swelling degree controls release rate
Water uptake allows protein to diffuse out
Swelling can be mediated by degradation
Controlled by physical or chemical erosion
Surface erosion mostly for physical gels
Constant release rate
Hydrolytical or enzymatic degradation
Protein size > mesh size
Stimuli lead to gel changes and drug release
Stimuli: temperature, pH, biomolecules, drugs, magnetic field, etc…
Adsorption of proteins through weak interaction (i.e. electrostatic, hydro-phobic, H-bonding)
Affinity with binding sites: heparin, anitbodies, chelating ions, peptides, molecular imprinting
Attachment of drugs to matrix via covalent bonds
Cleavable linker
Hydrolytical or enzymatic cleavage
Combination of multiple carriers: i.e.
protein-loaded micro/nanoparticles embedded in hydrogels
Release of multiple proteins through combination of mechanisms and rates
In-Situ Gelling Thermosensitive Hydrogels for Protein 111 (‘free-volume’) [116]. On the other hand, when the hydrogel pores are smaller than the protein diameter, swelling or erosion/degradation (bulk or surface) is needed for release. Generally speaking, the majority of the gel matrices reported to date exhibit diffusion-controlled release, following Higuchi’s kinetics, implying that the release is proportional to the square root of time [117]. This release profile was demonstrated to be particularly beneficial for the delivery of several growth factors for tissue engi-neering applications. For example, Brown et al. and Boerckel et al. demonstrated that bone regeneration was enhanced when rhBMP-2 was released down a concen-tration gradient [118,119]. The diffusional spatiotemporal growth factor presenta-tion proved effective not only for bone regenerapresenta-tion, but also for other engineered tissues like blood vessels [120,121]. The Ritger-Peppas equation is often used to fit release data and determine the underlying release mechanism: Mt/M∞=ktn with Mt/ M∞ the fractional drug release at time, t and k constants incorporating structural and geometric characteristics of the device. If n=0.5, the release is governed by Fickian diffusion. If n=1, molecules are released by surface erosion, while both mechanisms play a role if n has a value between 0.5 and 1 [122,123]. Swelling-controlled systems depend on water uptake and changes in drug diffusivity within the matrix. Swelling increases polymer flexibility and makes pores bigger, resulting in higher drug mobil-ity with n values that sometimes exceed 1. As a consequence, drug release depends on Fickian diffusion, polymer disentanglement and dissolution in water [124]. Erosion-controlled systems have increased in number since the development of synthetic bio-degradable polymers. In these systems, the mobility of the drug in the homogeneous non-degraded polymer matrix is limited and drug release is then governed by the degradation rate of the polymer, porosity increase, and drug diffusion mainly in the surface of the polymer matrix. The possibility to tailor the release kinetics of proteins from hydrogels can be achieved through the use of excipients or by changing the crosslinking density of the polymer network. In this particular instance, the use of synthetic polymers is extremely advantageous because they offer the opportunity to fine tune their chemical structure to achieve modular release. However, also natural polymer networks can be tailored to some extent by changing polymer concentration and crosslink density. Also, extensive work on the development of hybrid networks in which natural polymers are synthetically modified or combined with synthetic poly-mers has been proposed to combine beneficial properties of both kinds of polypoly-mers with respect to mechanical properties, tailorability in terms of biodegradability as well as tissue and cytocompatibility.
The geometry of the hydrogel-based depot also affects the release rate and the duration. Generally, the bigger the device, the longer the diffusion distances become and the longer the release consequently lasts. In recent years, patterning techniques utilized in tissue engineering (i.e., stereolithography, bioprinting) have been used to adjust surface area and consequently modulate release profiles of the biomolecules [94,125,126]. Deviation from the described release behaviors is observed when a spe-cific stimulus—typically temperature, pH, presence of certain molecules—leads to a physical or chemical change in the network structure; for instance, hydrogels swell or shrink in response to a certain trigger, thereby modulating the release of encapsulated drugs/proteins.
112 Handbook of Polymers for Pharmaceutical Technologies 4.3.1.2 Covalent Binding
Proteins, mostly by exploiting their reactive amine and thiol groups, can be covalently bound to polymer matrices via functional groups like hydroxy, amine and carboxyl groups that, if not naturally present in the structure of the polymer, have to be intro-duced by functionalization reactions, blending or copolymerization. The release of the protein is mediated either via hydrolysis or reduction reactions or (cell-mediated) enzymatic cleavage. This type of mechanism leads to on-demand release of loaded pro-teins according to the real physiological needs. However, stability and maintenance of the biological activity of the protein may represent a problem.
4.3.1.3 Dual/Multiple Delivery Systems
Biomedical hydrogels that can deliver multiple proteins in a multimodal mode and provide desirable pore structure and porosity to potentially encapsulate cells, have con-siderable potential as future therapeutic tools in medicine. The use of protein-loaded microspheres embedded in the hydrogel structure is a common approach to multi-modal protein delivery.
Microparticles of acidic and basic gelatin were used as carriers for the individual deliv-ery of two different growth factors and were embedded in a hydrogel matrix, composed of glycidyl methacrylated dextran (Dex-GMA)/gelatin. This hybrid system allowed the independent release of BMP-2 or IGF-1 that facilitated cell attachment, proliferation, metabolism, and osteoblastic differentiation of cells in a synergistic manner [127].
Fibrin hydrogels were complexed with heparin-functionalized PLGA nanoparticles for the release of VEGF for revascularization purposes [128].
A similar approach was adopted by Burdick et al., who formulated PLGA microspheres into PLA-PEG-PLA hydrogels for the delivery of multiple neutrophins with individual release rate [129]. Similarly, Biondi et al. used PLGA microspheres in collagen gels and collagen-HA semi-interpenetrating networks. The protein release kinetics and delivery onset strongly depended on the complex interplay between protein transport through the PLGA matrix and in the collagen-based release media, and water sequestration within the scaffold [130,131]. Physical hydrogel blends composed of hyaluronan (HA) and methyl cellulose (MC) were designed by Shoichet et al. for independent delivery of one or more therapeutic proteins, from 1 to 28 days, for the ultimate application of spinal cord injury repair. To achieve a diversity of release profiles, they exploited the combina-tion of fast diffusion-controlled release of dissolved solutes from the HAMC itself and slow drug release from PLGA particles dispersed within the gel [132]. Methacrylated hyaluronan networks encapsulating alginate microparticles enhanced mesenchymal stem cell chondrogenesis following delivery of TGF-β3 in vitro and in vivo [133].