Advanced Hydrogels as Exosome Delivery Systems for Osteogenic Differentiation of MSCs: Application in Bone Regeneration

Hydrogels are three-dimensional systems with high water content and hydrophilic polymer chains harboring special traits, including biocompatibility and elasticity, with the possibility of modification of their chemical properties. They are also able to mimic the extracellular matrix (ECM) and are capable of acting as a cell/tissue growth medium. These properties of hydrogels justify their broad use in biomedical research, from drug delivery to a vast range of applications in regenerative medicine. Their ability to encapsulate cells and increase the proliferation and retention of the cells is also of critical importance [24,25].

Hydrogel fibers consist of a fibrous texture with variable sizes from several nanometers to several microns. Hydrogel fibers are used in tissue engineering due to their high surface-to-volume ratio, rapid cell interaction, and immobility. Hydrogel fibers ensure cell viability and cell dispersion. The use of hydrogel fibers in bone regeneration has shown great potential in tissue engineering. Preparation of hydrogel fibers usually consists of two steps: a spinning process and a linking process. Generally, various types of spinning methods are used: electric spinning, microfluidic spinning, wet spinning, gel spinning, 3D printing technology, and hydrodynamic spinning, among which electric spinning and microfluidic spinning are the most common methodologies. Gelatin is obtained from the degradation of collagen and is an excellent polymer for the production of hydrogel fibers. This is due to its natural binding to the arginine–glycine–aspartic acid peptide (RGD). Compared with microbeads, hydrogel fibers can be injected at the defect site with a syringe and remain at the implant site for longer periods of time. Perez et al. co-delivered cobalt (Co) and BMP with well-tunable core shell hydrogel fiber scaffolds to induce angiogenesis and ossification in a rat calvarial defect [43]. On the other hand, hydrogel fibers have disadvantages such as poor mechanical strength and high swelling properties.

Poly vinyl alcohol (PVA) hydrogel fibers are highly swollen and therefore burst and release their components. To reduce the unwanted adverse effects due to the sudden release of the incorporated medication and growth factors, a research team successfully extended the duration of the drug by modifying the surface of the hydrogel fibers. Im et al. found that carbon–fluorine (c-f) bonding to the surface of PVA electrospun fibers through fluorination could significantly reduce the probability of a fiber reactive dye (Procion Blue) bursting and increased the diffusion time by 6.7-fold [44]. Furthermore, due to the stretched structure of hydrogel fibers, they lack favorable mechanical properties. Composite materials have been prepared from hydrogel fibers, such as a calcium phosphate cement (CPC)–hydrogel fiber structure, to improve the mechanical properties of hydrogels. Wang et al. synthesized a full-cell CPC hydrogel fiber composite material using wet spinning and mixing with a CPC paste. The mechanical properties and strength of this scaffold (8.5 ± 8.5 million Pascal) exceed that of spongy bone [45]. This is a promising result with application potential in the treatment of a wide range of bone defects in weight-bearing bones [39].

Traditional methods have been used for the preparation of microbeads. These technologies include emulsification, microfluidics, electrostatic droplet extrusion, coaxial air jetting, and in situ polymerization. However, traditional methods cannot provide uniform microbeads with a small size. The non-equilibrium microfluidic technique is now used for the preparation of smaller-size hydrogel beads (a size less than 100 µm) [46]. During the process of non-equilibrium microfluidics, polymer materials are inserted into a non-equilibrium water in oil (W/O) bonding that contains hydrogel molecules, in which the water molecules are dissolved into a continuous phase. The hydrogel precursors are condensed into the W/O droplets rapidly and form microbeads that are smaller than those formed under the conventional methods. Moshaverinia et al. have used injectable alginate hydrogel microbeads to encapsulate MSCs derived from dental tissue (gingival mesenchymal stem cells (GMSCs) and periodontal ligament stem cells (PDLSCs)). Mineralization of the inside of, and around, the microbeads was achieved and the cells remained viable post-implantation [47]. Furthermore, chitosan and collagen microbeads were used to encapsulate adult bone marrow stem cells via double cross-linking mechanisms by Wang and co-workers. The researchers showed that the expression of the transcription factor osterix (osx) and osteocalcin was increased and a significant deposition of bone minerals was achieved within the osteogenic medium [48].

Hydrogel nanoparticles (nanogels) are formed by physical or chemical cross-linking. They are a group of spherical nanoparticles capable of swelling in aqueous media. Nanogels have a great advantage in bone regeneration due to their good biocompatibility and desirable mechanical properties. Nanogels are typically synthesized by emulsion polymerization, such as reverse emulsion and distillation–precipitation polymerization, by rapidly stirring a solution at high temperature to disperse it steadily. In terms of their design and ease of preparation which has a wide range of polyvalent biological compositions, and their uniformity, adjustable size, stability, and high drug loading capacity, nanohydrogels are considered to be suitable systems for drug delivery. However, drugs loaded in nanogels are released quickly due to a difficulty with controlling the cross-linking point during gel formation. To solve this problem, a nanoscale structure needs to be designed to control the drug’s release [39]. Seo et al. have produced PEG nanogels with a diameter of less than 200 nanometers that turn into gel immediately after injection into the target site [49]. Yang et al. used PEG nanogels prepared by reverse microemulsion polymerization (IPMP) to carry a QK pro-angiogenic peptide. The cross-linking density of the nanogels was controlled by changing the mole fraction of the cross-linker to balance the release kinetics of the peptide QK (amino acids 17–25 of the VEGFA protein) [50]. Miahara et al. compared the use of a cholesterol-bearing pullulan (CHP) nanogel membrane with a collagen membrane in healing a parietal bone defect of an adult Wistar rat. They showed that the method increases the bone formation more effectively such that the newly formed bone in the nanogel group and the main bone are not histologically recognizable [51]. A study has applied acrylate-modulated CHP nanogels to deliver recombinant human fibroblast growth factor 18 (FGF-18) and recombinant human BMP-2 to defective bone. It was shown that the method is able to activate bone cells and eventually regenerate bone [52]. Although nanogels are suitable options for transporting proteins and growth factors to bone and inducing bone growth, the design of adjustable hydrogels that provide a high degree of mechanical stability and permanent release is essential to creating effective therapies for bone repair [39].

Several methods have been used for the fabrication and preparation of 3D scaffolds (Table 2). Conventional techniques such as blending, freeze-drying, salt leaching, and gas foaming have limited in terms of functional porosity, the requisite pore shape, geometry, and interconnectivity of scaffolds [41]. Such methodologies do not offer the native tissue’s cellular organization and totally rely on the manual seeding of the cellular components on prefabricated scaffold structures. Moreover, they are not cost-effective and their utilization is time-consuming and operator-dependent. Hence, different technologies, such as electrospinning and 3D (bio) printing, have been adopted for the fabrication of scaffolds with desired properties [53].

Electrospinning is considered to be a very common scaffold fabrication method. This technology enables us to create nanofibrous scaffolds with interconnected pores. In electrospinning, an external electric field is applied to draw charged threads of polymer melts or solutions as very thin stream jets from a capillary tube towards a collector plate. Fibers in the nanometer diameter range are produced and deposited continuously to create a scaffold. This process has the potential to incorporate various biomolecules and composite materials [55,56]. In electrospinning, different polymers of natural or synthetic origin can be used. The natural polymers used include collagen, gelatin, silk fibroin, chitosan, silk fibroin, and hyaluronic acid. The synthetic polymers that are often used in electrospinning include polyethylene oxide, poly (lactic acid) (PLA), poly (ɛ-caprolactone) (PCL), and copolymers, such as poly (lactic-co-glycolic acid) (PLGA) and poly (L-lactide-co-caprolactone). For all of these polymers, an electrospinning methodology has been used to produce scaffolds with desired properties for drug-delivery applications and specific tissue regeneration projects [57]. Several criteria should be considered when designing electrospun bone scaffolds. These include a suitable mechanical strength and tunable biodegradation kinetics, scaffold biocompatibility and biodegradability, the pore size, inter-related open porosity for growth factors, and most importantly a sterile environment for cell seeding and eventual cell growth [13]. Coaxial electrospinning is a novel technique for the preparation of core–shell nanofibers of a gelatin–chitosan generation. The resulting scaffold has a cationic nature on the surface (chitosan), while the gelatin in its interior highly favors the attachment of cells and their proliferation. As reported by Chen et al., deposition of HA onto the surface of the nanofibers results in increased mineralization efficiency and cell adhesion [58]. Bochicchio et al. used an electrospun composite scaffold composed of poly (D, L-lactic acid) (PDLLA) and gelatin with RKKP glass and ceramics embedded inside the nanofibers. Bioactive glasses like RKKP can stimulate the formation of a hydroxycarbonate apatite layer on the scaffold to improve their interaction with the bone surface in physiological conditions and also enhance osteoconductive properties [59].

Three-dimensional printing has several advantages over traditional fabrication methods. The 3D printing technology provides increased precision in designing the structure, which is required to support cell growth, proliferation, and migration. The technology enables the researcher to gain control over the pore size and shape, the level of porosity, and also the interconnectivity of pores. Highly controlled conditions for scaffold design, higher structural complexity, flexibility, and patient-specific demands are achievable [37]. The three-dimensional bioprinting process leads to printed scaffold structures with the architectural information of the defective tissue. The 3D structure of the required scaffold is obtained through instrumentational imaging that includes computerized tomography (CT) and magnetic resonance imaging (MRI). In addition, for the patient-specific and anatomically identical reproduction of the tissue, specific computer programs are used [37].

The major technological advancements utilized for the deposition and patterning of biological materials onto a scaffold are inkjet printing, laser-assisted printing, and micro-extrusion (Figure 1). Micro-extrusion technology is commonly used and consists of temperature-controlled material handling. It is important, however, to note that in micro-extrusion the cell viability is lower than in inkjet-based bioprinting. This is due to the fact that the cell survival rates are reduced with increasing extrusion pressure and at higher nozzle gauge values [63].

Bio-inks play a crucial role in the advancement of the technologies of functional organs and tissues in tissue engineering via 3D bioprinting. Bio-ink-related biomaterials contain collagen, agarose, alginate, gelatin-methacrylates, HA, chitin, silk, cellulose, and combinations of these materials. Nevertheless, there are major challenges to the use of bio-inks. These challenges include cell encapsulation, minimal conditions for bioprintability and cytotoxicity, high morphologic stability after printing, and biocompatibility. The designed scaffold must be able to keep its shape under wet conditions to be able to support cell proliferation and adhesion. Noh and co-workers used GelMA, HA, and hydroxyethyl acrylate (HEA) to promote the hydrogel’s mechanical stability and cellular interaction [54]. Biological factors that are involved in communication between hydrogels and stem cells include stem cell survival, polymer types, stiffness, porosity, degradation, and compatibility [64].

Strategies for bone tissue engineering provide novel approaches to bone repair through recapitulation of the developmental process of endochondral ossification. Based on previous studies, a 3D-printing modality could be exploited to engineer cartilage tissues with complex geometries and precisely controlled internal architectures to support vascularization during endochondral bone repair. Pluronic ink as a thermo-responsive hydrogel was used by Daly and coworkers to devise networks of interconnected microchannels inside MSC-laden GelMA hydrogels and printed a range of different microchannel diameters in arbitrary complex geometries [65].

Almost all scaffold fabrication methodologies lack the ability to provide the heterogeneous multiphasic porous architecture of the target tissues. Although it is possible with electrospinning technology to produce a dense network of micro/nanofibers with pore sizes and fiber diameters that closely resemble the architecture of native ECM, this technology lacks the potential to generate the expected three-dimensional structures of relevant functional tissue. On the other hand, 3D bioprinting provides scaffolds with higher dimensional control and better reproducibility. However, using this technique results in scaffolds with thicker fibers and larger pore sizes compared with electrospinning [66]. Mellor et al. used a combination of electrospinning and 3D bioprinting techniques. Using a bio-ink consisting of human-adipose-derived stem cells (hASC), they developed a reproducible and simple scaffold that incorporated both a micro-scale and a nano-scale fibrous architecture mimicking the heterogenous tissue structures. They showed that hASCs were adherent to and proliferated in the cultures, both in 3D-bioplotted and electrospun scaffolds. They showed that there was a minimal number of dead cells after 21 days in the tissue culture and that this combined scaffold structure could be easily implanted into an osteochondral defect without any problems (breaking or delamination of the scaffold) [67]. Therefore, the incorporation of electrospun nanofiber segments into bioactive materials with 3D-printed scaffolds improved cell adhesion and proliferation [68].

MSC-derived exosomes have been exploited for different therapeutic approaches to vascular diseases such as stroke (cerebral infarction), cardiac fibrosis, and liver ischemia/reperfusion injury [86]. It has been shown that exosomes with enriched miRNA content could be essential players in intercellular communication in the nervous system. For example, through miR-132, exosomes may become involved in cell-to-cell signaling and mediate neural regulation of brain vascular integrity [87]. It has been shown that exosomes derived from human umbilical cord MSCs (hucMSC exosomes), human adipocyte MSCs (haMSC exosomes), and human induced pluripotent stem-cell-derived MSCs (hiPS-MSC exosomes) promote skin wound healing by delivering various functional soluble signaling molecules, such as cytokines, RNAs, and proteins [88]. By exploitation of exosomes, the cancer cell secretome modulates stromal cell fate and induces the differentiation fibroblast cells into myofibroblast lineages.

Exosomes may trigger the differentiation of stem cells without the help of growth factors and cytokines. Wang and co-workers obtained exosomes from undifferentiated human mesenchymal stem cells (hMSCs) and compared their miRNA contents after osteogenic lineage differentiation [89]. The expression levels of miRNAs involved in osteogenic differentiation, such as miR-181a, miR-221, miR-885-5p, miR-155, and miR-320c, were lower in exosomes isolated from differentiated cells, while in undifferentiated cells the expression levels of miR-148a, miR-199b, miR-218, miR-135b, let-7a, miR-219, miR-203, miR-299-5p, and miR-302b were increased [89]. Many investigations have indicated that functional miRNAs may be selectively loaded into exosomes, which in turn serve as important regulators of angiogenesis and other endothelial cell functions [90]. The microRNA miR-885-5p negatively regulates the osteogenic differentiation of bone-marrow-derived MSCs by inhibiting the functions of bone morphogenetic protein 2 (BMP2). Furthermore, the suppressive effects of miR-885-5p on osteogenic differentiation involve the Wnt pathway as they result in downregulation of Wnt5a mRNA [91]. On the other hand, let-7a positively regulates osteogenesis lineage differentiation by influencing high mobility group AT-hook 2 (HMGA2) and blocking the adipogenic differentiation of MSCs [92]. In a similar way, miR-218 positively regulates osteogenic differentiation. This effect has an impact on the Wnt/ß-catenin pathway, which in turn is an important player in the osteogenesis of adipose-derived stem cells [93]. The stability and structural retention of exosomes are very important for their useful application (Figure 3).

Chitosan hydrogels have high biocompatibility, mimic the natural extracellular matrix, and facilitate cell migration, adhesion, and proliferation. Zhang et al. used thermosensitive chitosan hydrogels to improve the in vivo retention and stability of exosomes as well as enhance their therapeutic capacity [94]. In particular, 3D-printed PLA loaded with human gingival MSCs (hGMSCs) and EVs was shown by a MicroCT assessment to enhance bone regeneration and the vascularization process [95]. Table 3 summarizes the therapeutic effects of loading osteogenic miRNAs into exosomes in pre-clinical animal model systems.

Data are contained within the article.