Nanostructured injectable cell microcarriers for tissue regeneration

Abstract

Biodegradable polymer microspheres have emerged as cell carriers for the regeneration and repair of irregularly shaped tissue defects due to their injectability, controllable biodegradability and capacity for drug incorporation and release. Notably, recent advances in nanotechnology allowed the manipulation of the physical and chemical properties of the microspheres at the nanoscale, creating nanostructured microspheres mimicking the composition and/or structure of natural extracellular matrix. These nanostructured microspheres, including nanocomposite microspheres and nanofibrous microspheres, have been employed as cell carriers for tissue regeneration. They enhance cell attachment and proliferation, promote positive cell-carrier interactions and facilitate stem cell differentiation for target tissue regeneration. This review highlights the recent advances in nanostructured microspheres that are employed as injectable, biomimetic and cell-instructive cell carriers.

In tissue engineering, scaffolds play a critical role in guiding cell adhesion, behavior and function. Implantable 3D porous scaffolds can serve as supportive structures for cell adhesion and provide the necessary spatial and physical support for cellular activity [1–3] (Figure 1A). However, the repair of irregularly shaped tissues or defects is often hard to achieve with a prefabricated scaffold, or at a minimum requires complex design and fabrication of a one-off scaffold to achieve accurate fit [4–7]. An injectable scaffold designed to fill any defect site no matter the geometry could therefore be advantageous, simplifying the scaffold design and surgical procedure [8,9]. Injectable scaffolds also could lead to shortened recovery times thanks to the minimal invasiveness of an injection versus the surgical implantation of a prefabricated scaffold. Cell carriers, already used for growing large numbers of cells in the pharmaceutical industry, have increasingly found a role in tissue engineering scaffold design due to their injectability. Various new injectable carriers have been developed to deliver cells and regenerate irregularly shaped tissue defects [10–12]. Like other scaffolds used in tissue engineering, cell microcarriers have been designed to deliver growth factors and cells for tissue regeneration [13–16] (Figure 1B & C) and have been constructed out of a wide variety of materials, each with specific advantages and disadvantages. In addition to growth factor/small molecule incorporation, the nanoarchitecture of the injectable particle has also been discovered to greatly influence cell behavior and tissue regeneration. This paper reviews a new generation of nanostructured microspheres developed as injectable cell and drug carriers for tissue regeneration. We first briefly summarize the basic design goals and strategies for 3D implantable nanostructured scaffold fabrication, with a specific focus on the importance of biomimicry and biodegradability. We then review various biodegradable microspheres for cell delivery with a special focus on drug delivery, nanostructured architectures and functionalization. This review concludes with the authors’ perspectives on current challenges and future trends of microsphere-based cell delivery for tissue regeneration.

Figure 1. . A schematic showing the three typical strategies for cell delivery and tissue regeneration.

(A) Cell-seeded 3D porous scaffolds for implantation. (B) Cell-mixed hydrogel prepolymer solution for injection to the targeted site for in situ gelation. (C) Cell-laden microspheres, with or without in vitro culture, for injection to the targeted site.

Tissue engineering scaffold design

An ideal scaffold should perform the structural and biochemical functions of the extracellular matrix (ECM) until the cells interacting with them produce their own natural ECM and ultimately integrate into the surrounding tissue [17]. In the interim, a scaffold should provide an environment with similar mechanical, geometrical and topological cues for the native tissue cells to interact with. The chemical and physical features of a general 3D scaffold can be designed to mimic those of the ECM in order to facilitate and enhance cell adhesion, proliferation, migration and differentiation. In addition, scaffolds should be biodegradable to enable cell modeling/remodeling and neotissue formation without leaving behind residual materials afterward. Porous biodegradable polymers are therefore widely utilized in tissue engineering, thanks to their customizability with respect to biodegradability, biomimicry, porosity, mechanical properties and ability to maintain a predesigned tissue structure.

Biodegradable polymers for scaffolding can be categorized into naturally derived polymers and synthetic polymers. Naturally derived polymers, for example, polypeptides and polysaccharides, have the advantage of biological recognition that might support cell development. However, there are also disadvantages associated with their use. Collagen, for instance, is the main component of ECM and offers an ideal environment for cell growth and differentiation [18], but there are concerns regarding the immune response and potential disease transmission from collagen extracted from animals [19]. As a result, there is an advantage to using recombinant collagen or gelatin over xenogenic sources [20,21]. As an alternative, synthetic biodegradable polymers have become widely used as scaffolding materials, because of their design flexibility in composition and easy fabrication of porous structure to achieve desired functions. Poly(lactic acid), poly(glycolic acid) and their copolymer poly(lactic-co-glycolic acid) (PLGA) are among the few US FDA-approved synthetic materials for certain human clinical products (e.g., degradable sutures, stents, screws, bone plates and wound dressings). Critically, the degradation rate of these poly(α-hydroxy acids) can be controlled by varying polymer compositions, with the degradation products eliminated through metabolic pathways. For these reasons, much research has focused on the fabrication of 3D scaffolds from poly(α-hydroxy acids) [22–31].

Because these synthetic polymers enable flexible fabrication into various chemical and physical structures, various material processing strategies have been developed to create tissue engineering scaffolds with characteristics of the natural ECM [16,32]. In particular, a topological environment dominated by nanofibers is an important nanosized physical feature. These fibers resemble the fibrillar structure of collagen in both morphology and size, along with many other ECM components. Various synthetic nanofibers have been found to enhance cell–matrix interactions, resulting in improved cell adhesion, proliferation and differentiation [33–41]. Self-assembly, electrospinning and phase separation are the three techniques currently used for the fabrication of macroscopically sized scaffolds featuring what we define as ‘nanofibrous (NF) morphology’. Among these techniques, thermally induced phase separation (TIPS) cannot only generate synthetic nanofibers, but also can be combined with various fabrication techniques to generate a predesigned micropore network and 3D geometry [2,5,7,42–45]. These 3D porous NF scaffolds facilitate and enhance stem cell delivery for the regeneration of various tissues [46].

As discussed previously, one downside to prefabricated 3D scaffolds is the necessarily invasive procedures required for their implantation into a patient (e.g., surgical incision and subsequent trauma and comorbidity). Thus, there have been a variety of injectable scaffolds and materials developed using both natural and synthetic materials to address this problem. Hydrogels with both natural and synthetic origins have been used as injectable cell carriers, due to their biocompatibility and gel-like mechanical properties that are amenable to injection [8,47–55]. A suspension of prepolymer/macromer and cells can be injected and polymerized to form a gel in situ, achieving an accurate fit to the site where tissue regeneration is desired (Figure 1B). Hydrogel formation occurs after crosslinking macromers such as by using chemical crosslinkers [56] and radiation [57]. The resulting hydrogel network becomes highly swollen with tissue-like high water content. While hydrogels can encapsulate cells and bioactive molecules, hydrogels are not currently used for regeneration in clinics possibly due to several shortcomings. Most hydrogels provide insufficient cell anchorage sites, which are necessary for the viability of anchorage-dependent cells [58,59]. In addition, during in situ hydrogel formation, cell mobility and cell–cell interactions are restricted, as well as the host–implant integration at the cellular level as the cells are effectively ‘caged’ away from one another by the hydrogel. Various strategies have been employed to solve these problems [60–64]. For example, microchannels can be created by selectively shining light on a photodegradable hydrogel, allowing for subsequent migration of the encapsulated cells [60]. Tissue proteins can be incorporated into the hydrogel network to improve cell–matrix interactions [63]. Hydrogels with cell-mediated degradation can also be designed by incorporating peptide-based linkages that are susceptible to matrix metalloproteinases – these hydrogels allow for superior cell invasion/migration and thus better tissue integration [65]. There are newly developed advanced injectable (shear-thinning) and self-integrating hydrogels, which can be preformed into cell-loaded 3D gels and remain injectable into the defect to instantly recover their gel status upon the completion of injection [66]. Such dynamic hydrogels also impart cells higher mobility and easy integration with the host tissue.

Nevertheless, the various shortcomings of conventional hydrogels have inspired biomedical engineers to turn to alternative strategies for creating an injectable tissue engineering platform. Cell carriers such as microspheres largely obviate the weaknesses displayed by hydrogels, as they allow for cells to directly interact with each other and with any surrounding tissue. Here, we define a nanostructured injectable cell carrier as a particle with architectural features (e.g., fibers, channels, domains etc.) on the order of nanometers. These cell carriers consequently can be incubated with cells for delivery via syringe. Much like any other tissue engineering scaffold, proper selection of the base material and processing technique in order to optimize biodegradability, cell adhesion and other important properties is of critical importance.

Injectable microspheres for tissue engineering

In order to optimize microspheres for tissue engineering and regeneration, several design requirements need to be satisfied. First, the microspheres should be both biodegradable and biocompatible. Second, the degradation rate should closely match the rate of neotissue formation, and demonstrate appropriate longevity. Third, the size and morphology of a microcarrier should be suitable to carry cells. Fourth, a microsphere's chemical composition and surface architecture should facilitate cell adhesion, proliferation and differentiation. In addition, bioactive and cell-instructive microspheres which can direct stem cell differentiation, phenotype maintenance and facilitate target tissue regeneration may be necessary. These ‘cell-instructive’ elements may be built into the microsphere through its architecture or mechanical properties, or could be incorporated through the attachment of growth factors or cytokines.

Biodegradable microspheres have been used as cell carriers with injectability, controllable biodegradability and capacity for drug incorporation [67]. Compared with hydrogel-based injectable carriers, microspheres could provide sufficient anchorages and better facilitate cell attachment for anchorage-dependent cells. Microspheres were originally employed as a cell culture system to produce biological cell products [68]. The de novo production of extracellular matrix by cells seeded on particulate microcarriers resembled many features of the tissue of origin [12], which inspired researchers to explore the potential of using microspheres as cell delivery vehicles for tissue engineering. These traditional solid microspheres, however, are often nonbiodegradable and lack biomimetic surface structure to interact with cells.

Optimizing cell attachment and adhesion is another important factor in microsphere design. Typically, cells are directly mixed with the microspheres in a suspension culture and injected into the defect site with or without preculture in differentiation media (Figure 1C). To enhance cell attachment, microspheres can be precoated with a variety of proteins (such as fibronectin), adhesive factors or even a solution of serum. If the attached cells are stem cells, the microsphere/cell construct can then be cultured in vitro in any desired media cocktail for an optimized length of time to induce differentiation prior to injection in vivo, retaining injectability. Porosity is another factor in design, permitting efficient nutrient/waste transfer and therefore a positive microenvironment for cell function [9].

Cell–microsphere interactions (whether physical or chemical) are critical for the resulting phenotype and performance of the cell/microsphere construct. Understanding these interactions can guide engineers to better microsphere designs. Thus, to enhance cell–particle interactions, researchers are now developing technologies to tune the physical and chemical structure of the cell carriers, with a particular focus on mimicking the natural ECM environment of the cells [16,69]. On the one hand, scaffolds with a nanoarchitecture resembling the natural ECM nanostructure can enhance cell–matrix interactions and tissue regeneration [9,46]. For example, by mimicking the architecture of natural collagen fibrils, NF scaffolds enhanced multiple types of tissue formation, including bone [46], dentin [70], intervertebral disks [71], cartilage [9,72] and blood vessels [39]. In another example, inorganic/organic nanocomposites which mimic mineralized tissue ECM composition at the nanoscale have been used to create cell carriers for hard tissue regeneration [72–77]. On the other hand, engineering scaffolds to bind and present appropriate biochemical signals (e.g., growth factors/cytokines/etc) is an important strategy for regulating stem cells in tissue regeneration [78]. For example, ECM-bound growth factors are important regulatory signals, which are widely incorporated in biomaterials for morphogenesis in tissue engineering through various delivery/presentation strategies [16,46]. Therefore, taking into consideration both the physical architecture and the incorporation of specific signaling moieties into the design of injectable microspheres is critical for their function and optimization.

Similar to 3D monolithic scaffolding, both natural and synthetic polymers are used to fabricate biodegradable microspheres. One such example is Cellagen^®, a cross-linked type I collagen microcarrier, which can support chondrocyte proliferation and phenotype expression [79]. The chondrocyte proliferation rate on Cellagen was more than 20-times higher than that of cells cultured in monolayer, with abundant type II collagen secretion and thus enhanced expression of the chondrocytic phenotype. These collagen microspheres were also found to support osteoblast proliferation and the production of bone matrix proteins [80]. Unfortunately, collagen has relatively poor mechanical properties. In one reported approach to address this issue, a collagen/mesenchymal stem cell (MSC) suspension was fabricated into microspheres, where the cell-induced contraction improved mechanical stability [81]. The MSCs were also shown to retain viability and multipotency to differentiate down both chondrogenic [82] and osteogenic [83] lineages.

Alternatively, gelatin has a chemical structure similar to collagen (denatured from collagen), while its mechanical properties and degradation rate can be controlled through adjusting the cross-linking densities. There is also a lower level of concern over provoking an immune response or disease transmission with gelatin when compared with collagen. Macroporous gelatin CultiSpher G^® microspheres were used as a cell carrier for human nasal chondrocytes to form neocartilage, which were subsequently seeded on 3D scaffolds and implanted in nude mice subcutaneously [84]. Significantly, more proteoglycans were deposited in the mice implanted with chondrocytes grown on microspheres over those grown on 2D flasks. In addition, gelatin can be positively or negatively charged at neutral pH, allowing electrostatic interactions to take place between a charged biomolecule and gelatin of the opposite charge, forming polyion complexes [85]. These charged biomolecules could be released to promote tissue regeneration.

Chitosan, the partially deacetylated derivative of chitin, is another widely used natural polymer for biomedical applications [86]. Chitosan is biodegradable and has good mechanical properties. In addition to being used extensively for creating monolithic scaffolds, chitosan has also been fabricated into microspheres for potential cell delivery and tissue regeneration purposes [87–90]. The advantages of chitosan include biocompatibility, availability in large quantities and its intrinsic antibacterial properties [86]. Due to its structural similarities to glycosaminoglycans (GAGs), which are important components of liver, cartilage and other tissue ECMs, chitosan has also been tested as a microcarrier material for hepatocyte delivery. These chitosan microcarriers facilitate cell aggregate formation and proper liver phenotype expression, showing promise for the development of a hybrid bioartificial liver support system [91]. Similar to gelatin, chitosan can also form polyion complexes with charged proteins or drugs, creating a bioactive agent delivery system capable of slow release [92]. This combination of drug delivery and cell delivery within one microcarrier system will be discussed in greater detail later in this review. Chitosan has also been combined with aldehyded 1-amino-3,3-diethoxy-propane–hyaluronic acid to form an injectable hydrogel for skin tissue engineering [93].

Alginate microspheres are a widely studied cell carrier for tissue engineering [94–98]. Cells which are more sensitive to their environment can be successfully encapsulated in alginate thanks to its mild processing conditions (in the presence of Ca²⁺) [99,100]. In addition, alginate has been used for bone regeneration, considering its calcification capacity in vivo. This results from interactions between Ca²⁺ in the cross-linked alginate and the surrounding phosphate ions [101], making it a useful base material for orthopedic applications. However, the slow and uncontrolled degradation in vivo and poor cell adhesion properties of alginate must be addressed [102].

Composite microspheres can also be fabricated from natural polymers, combining the strengths of individual materials to overcome their respective weaknesses. For example, to overcome its relatively poor mechanical properties, collagen can be cross-linked [79–80,103–104] or blended with other polymers, such as agarose [105] and chitosan [106], resulting in hybrid microspheres with improved mechanical stability. Porous chitosan microcarriers have been coated with collagen to enhance cell attachment for cartilage regeneration [107]. Alternatively, minerals such as hydroxyapatite (HAP) can be coated onto natural polymers to create better scaffolds for bone tissue engineering. In one study, mineralized chitosan microspheres were used as bone fillers and found to support MSC attachment [87].

Because natural polymers can provoke an immune response, and could transmit undetected pathogens, synthetic polymers receive considerable attention as microsphere materials. A wide variety of synthetic microspheres for tissue engineering have been developed, and many show potential for enhancing tissue regeneration [9,108–110]. For example, chondrocytes were attached and cultured on PLGA microspheres in vitro in a cylindrical mold for 8 weeks, successfully maintaining their phenotype. The construct stained positive for type II collagen and had higher tissue mass and GAG content versus chondrocytes cultured directly on the mold without microspheres [25]. In another study, rabbit chondrocytes were mixed with the PLGA microspheres and immediately injected subcutaneously in mice. Cartilage formed after 9 weeks, as shown by positive GAG and type II collagen staining. By comparison, bare chondrocyte transplantation without microspheres and microsphere injection without chondrocytes led to significantly less cartilage formation [111,112]. Thus, the combination of microspheres with cells ex vivo prior to implantation can lead to enhanced tissue regeneration. Other polymers could also be blended with poly(α-hydroxy acids) to fabricate microcarriers for cells [67,113–114]. Other synthetic injectable platforms include polycaprolactone particles featuring cell-adhesive polydopamine for human neural stem cell attachment [115].

While PLA, PGA and PLGA have been extensively used in bone and cartilage tissue engineering, synthetic polymers often lack biological activity such as osteoconductivity and osteoinductivity for bone regeneration. In contrast, calcium phosphate (CaP) or bioactive ceramics/glasses have been developed as bone fillers due to their excellent biological properties and compressive mechanical properties [116], which could simulate the inorganic phase of bone. CaP cements are popular injectable bone fillers, but have a slow rate of resorption and require the generation of a macroporosity to allow cell migration and ingrowth [117]. In addition, CaP cements are unable to carry cells due to the poor chemical environment. Thus, for direct cell delivery to bone defect sites, inorganic microspheres are widely studied [118–120]. As one example, calcium carbonate microspheres were found to stimulate mineralization and MSC differentiation in vitro and enhanced bone formation in vivo [121]. However, the fabrication of inorganic matters into a spherical shape is challenging, and often requires the incorporation of a polymer phase that subsequently needs to be removed after the spheres formation [67].

Alternatively, inorganic components can be dispersed or blended within polymers to fabricate composite microspheres, taking advantage of the processability of polymers and the osteoconductivity of inorganic materials [122–125]. Importantly, the polymer phase can be biodegradable with a tunable degradation rate (achieved by varying the chemical composition or the molecular weight of the polymer), while also taking advantage of the osteoconducive properties of the inorganic phase present. Additionally, drugs or growth factors could be loaded to implement drug delivery.

Like natural polymers, synthetic polymers can also be mixed with other materials to create composites which address shortcomings or enhance properties of the pure polymer. Incorporating HAP, an important inorganic component of bone ECM, into synthetic microspheres to fabricate a composite material with enhanced osteoconductivity is one example of a composite microsphere. The nanoscale organization of HAP within a polymeric matrix provides increased hydrophilicity, wettability, surface area and surface roughness, which leads to enhanced osteoblast adhesion and osteoconductivity and better bonding to host bone for long-term functionality [126,127]. One approach for organizing HAP at the nanoscale is through in situ precipitation of apatite crystals within a polymer or macromolecular solution [122,123]. The resulting HAP-gelatin nanocomposite microspheres demonstrated significantly better osteoblast attachment, proliferation and phenotype expression in vitro, as compared with their microscale composite counterparts [122]. Nanosized HAP coating through chemical bonding is another strategy for building nanocomposites [128]. HAP-PLLA nanocomposite microcarriers generated with this technique were found to enhance bone marrow mononuclear cell adhesion and promote the subsequent clustering of MSCs and bone formation [129]. Pickering emulsification can also be used to build a nanosurface directly onto microspheres. Nanosized HAP is used as surfactant during emulsification to stabilize the oil–water interface, allowing it to then deposit onto the surface of the formed microspheres upon solidification [130]. Through adjusting the HAP content in water, the resulting HAP coverage on the surface of microspheres could be adjusted.

Recently, an electrodeposition technique was developed that resulted in a fast and high-quality mineralized coating on various substrates [131]. Control over the surface topography and Ca/P ratio was achieved through varying parameters like temperature and voltage (Figure 2). Electrodeposition was successfully performed on PLLA nanofibers, showcasing this technique's potential and flexibility in the mineralization of other substrates.

Figure 2. . Scanning electron microscopy (SEM) images of calcium phosphate deposition on electrospun PLLA nanofibrous scaffolds.

Electrodeposition was performed for 60 min at 60°C and (A) 2, (B) 3 and (C) 5 V.

Adapted with permission from He et al. [131]. © John Wiley & Sons, Inc. (2010).

Drug releasing microspheres as cell carriers

Molecular signaling plays a critical role in tissue regeneration. However, endogenous signaling molecules are often insufficient in type and/or quantity for the repair of critical-sized defects. Therefore, the delivery and sustained release of biological molecules could enhance tissue regeneration. Due to the sensitivity of cells to high concentrations of signaling molecules coupled with their short half-lives, a delivery vehicle is often required to safely transport them in vivo to where they are needed. Drugs and signaling molecules can be encapsulated in microspheres, which protect the bioactivity of various therapeutic agents. Biodegradable polyesters such as PLLA and PLGA are often used to fabricate micro- and nanospheres for drug encapsulation (Figure 3A), with the drug release profile determined by both the drug diffusion through the microspheres and degradation patterns of the polymers themselves. Therefore, this microsphere system is intended to leverage the dual functions of drug and cell delivery for tissue regeneration [26,132–137]. In one study, NGF-encapsulated PLGA microspheres were exploited as cell carriers for fetal rat (E16-E17) brain cells, generating neotissues with enhanced neuronal choline acetyltransferase activity in vitro. Upon transplantation into mice brains, the neotissues produced significant elevations in NGF content and NGF-induced biological activity over the course of 21 days [136].

Figure 3. . A schematic showing the implementation of drug delivery capability into cell microcarriers.

(A) Drug is directly loaded into microspheres through encapsulation/dispersion; (B) Drug is loaded on the microsphere surface through LbL process; (C) Drug-loaded nanospheres are encapsulated/dispersed into microspheres; (D) Drug-loaded nanospheres are deposited on the microsphere surface through LbL process.

LbL: Layer-by-layer.

Combining drug and cell delivery on the same microsphere simplifies designing a platform for tissue regeneration. However, it is a challenge to simultaneously optimize a microsphere degradation pattern to achieve a desired drug release profile, while at the same time ensuring that the microspheres provide sufficient longevity and mechanical support for cells until the neotissue is formed. To address this issue, one strategy would be to decouple the surface (in which the drug may be encapsulated) from the bulk material, allowing for tunability of both with a broad degree of independence from each other. The electrostatic layer-by-layer (LbL) technique can assemble programmable drug delivery capability onto cell microcarriers without changing the bulk properties of the microcarriers (Figure 3B) [138]. LbL deposition is done by alternating a coating of oppositely charged materials, with wash steps in between. A primary advantage of the LbL self-assembly technique is the generation of ordered structure at nanometer thickness (with an achievable resolution of 1 nm) on substrates of various shapes and sizes [139]. The drug load can be easily tuned with the number of layers incorporated [140]. The LbL technique has been used to encapsulate drugs onto microsphere substrates [141–143]. In one example, a multilayered polyelectrolyte delivery system was constructed on the surface of PLGA microspheres using LbL [138]. The layers incorporated heparin to facilitate the loading of bFGF with increased stability. The growth factor release kinetics was manipulated by changing the number of layers and amount of drug loaded, while retaining the bioactivity of the encapsulated growth factor. bFGF was released in a sustained manner and stimulated significantly higher mouse fibroblast cell proliferation in vitro compared with the direct addition of bFGF to the cell media.

In another strategy designed to decouple the bulk material of the microspheres from the material encapsulating the drug, drug-loaded nanoparticles could be encapsulated/dispersed within the microspheres (Figure 3C) [144,145], or assembled onto the surface of the microcarriers through the LbL technique (Figure 3D) [146–149] to introduce drug release capacity. In one study, different growth factors (TGF-β3, BMP-7, IGF and bFGF) were loaded into PLGA nanospheres and then attached to the surface of PLGA microspheres via the LbL technique [147]. MSCs were then cultured and attached to these structures, and the nanoparticle/microsphere constructs were able to induce MSCs to undergo chondrogenic, osteogenic and adipogenic differentiation both in vitro and in vivo (depending on the growth factor/s encapsulated) [147]. In a recent report, alginate was used to encapsulate drug-loaded nanoparticles and cells simultaneously to create composite microspheres, which enabled both protection and dual delivery of bioactive molecules and cells for tissue regeneration [145].

Nanofibrous microspheres for cell delivery

Similar to 3D monolithic scaffolds, it will be advantageous to impart nanofibers into microspheres to enhance cell–matrix interactions. However, it is challenging to create a microsphere with NF architecture using existing methods. Only recently, a novel technique has been developed to combine TIPS with emulsification to generate NF microspheres [9]. Linear PLLA was dissolved in tetrahydrofuran and emulsified into liquid microspheres in glycerol. The emulsion was then subjected to phase separation in liquid nitrogen, solvent extraction with ice water and freeze drying, leading to the formation of nanofibrous microspheres (NF-MS) (Figure 4A). NF-MS, composed entirely of nanofibers, has a much higher porosity and surface area as compared with traditional solid interior microspheres (SI-MS) [9]. A higher porosity provides more space for cell growth, ECM deposition and nutrient/waste diffusion and minimizes degradation products [16]. In addition, the higher porosity and higher surface area of PLLA NF-MS contributes to a faster degradation rate when compared with PLLA SI-MS.

Figure 4. . A schematic showing the fabrication of PLLA nanofibrous microspheres.

(A) Linear PLLA was employed to fabricate nanofibrous microspheres via emulsification and thermally induced phase separation techniques: a) SEM image of a representative nanofibrous microsphere; b) A high-magnification image of the microsphere, showing the nanofibers with an average diameter of about 160 nm. (B) Star-shaped PLLA was synthesized to fabricate nanofibrous hollow microspheres through emulsification, thermally induced phase separation and self-assembly. Dendrimer PAMAM (G2) was employed as an initiator for the synthesis of star-shaped PLLA. The colors show the successive PAMAM generations. c) SEM image of a representative nanofibrous hollow microsphere, showing a hole of approximately 20 μm on the microsphere shell. d) A 2D cross-section confocal image of the nanofibrous hollow microspheres, confirming the hollow structure of the microspheres.

Adapted with permission from Liu et al. [9]. © Nature Publishing Group (2011).

While the NF structure results in a microsphere with a high porosity (92%), it has since been demonstrated that a hollow structure can be generated inside the NF-MS to achieve an even higher porosity. This is realized through molecular self-assembly during the emulsification and phase separation, which requires no surfactant addition. Specifically, a new star-shaped polymer was synthesized. A dendrimer was used to initiate the ring-opening polymerization of L-lactide, giving rise to a star-shaped PLLA (SS-PLLA). The resulting polymer was subjected to emulsification and phase separation processes, leading to the assembly of nanofibrous hollow microspheres (NF-HMS) (Figure 4B). The hollow structure increases the porosity to 97%, which can be even further increased by decreasing the polymer solution concentration. The higher porosity and lower molecular weight led to an increased degradation rate and reduced the concentration of degradation products. Importantly, the hollow structure of the NF-HMS could accommodate more cell-produced ECM, resulting in more uniform and continuous tissue formation. The open pore size can be controlled (in the range of 10–50 μm) through changing the polymer solution concentration or blending linear PLLA with SS-PLLA during microsphere fabrication. A lower SS-PLLA concentration results in an increased open hole size, while blending linear PLLA with SS-PLLA leads to a decreased open hole size. In addition, the number of open holes increases as the size of the microspheres increases. Controlling the number and size of open pores can potentially be used to regulate both cell migration and morphology within the pores.

NF-HMS were evaluated as an injectable chondrocyte-carrying scaffold for cartilage regeneration [9]. Compared with solid microspheres, NF-HMS significantly enhanced cell proliferation, chondrogenic gene expression and cartilaginous matrix production (GAG, collagen type II and aggrecan). In a full-thickness rabbit knee osteochondral defect repair model, NF-HMS carrying chondrocytes repaired a critical-sized defect with a substantially thicker layer of cartilage and substantially higher GAG content than those in a chondrocytes-loaded PEG hydrogel, chondrocytes alone or the sham control groups 8 weeks after the injections (Figure 5A–D). In addition, NF-HMS/chondrocytes achieved superior integration with the underlying bone (Figure 5E–H), illustrating their advantage over the control carriers as an injectable tissue-engineering platform. Importantly, NF-HMS largely degraded after 8 weeks due to their hollow and NF structure, leaving regenerated high-quality cartilage in the defect, whereas conventional solid microspheres took a significantly longer time to break down. Furthermore, the neocartilage formed from a subcutaneous injection of chondrocyte-seeded NF-HMS for 8 weeks resulted in an aggregate modulus that matched native rabbit cartilage [9]. These results indicate that the NF-HMS are an excellent injectable cell carrier for cartilage regeneration and may also be an advantageous injectable cell carrier for regenerating other tissues.

Figure 5. . Evaluation of critical-size rabbit osteochondral defect repair 8 weeks after injection.

Safranin-O staining of: (A & E) NF-HMS/chondrocytes, (B & F) PEG/chondrocytes group, (C & G) chondrocytes-alone group and (D & H) sham control group, with increasing magnifications of 40× (A–D) and 200× (E–H). NF-HMS/chondrocytes group showed strongest Safranin-O staining (A & E) with thicker cartilage repair (A) and excellent cartilage-bone integration (A), compared with other groups, as well as a rounded cell morphology (E). (I) Aggregate modulus of PEG, PEG/chondrocyte, NF-HMS/chondrocyte and native rabbit cartilage groups. The aggregate modulus of NF-HMS/chondrocyte group was not statistically different from that of the native rabbit cartilage and was significantly higher than those of PEG/chondrocyte and PEG groups.

**p < 0.01.

Adapted with permission from Liu et al. [9]. © Nature Publishing Group (2011).

In the engineering of many tissue types, especially where cell–cell interactions are critical, scaffolds with an interconnected microporous structure are highly preferred. No technique existed which could impart both NF and interconnected microporous structures into microspheres until a recent study pioneered a method for doing so through the self-assembly of novel biodegradable polymers [150]. In this study, the effects of the SS-PLLA's molecular structure and functionalities on their self-assembly in TIPS and emulsification processes were thoroughly investigated. It was discovered that the polymer arm length determined the nanoscale self-assembly structure, while the density of functional hydroxyl groups determined the microscale self-assembly behaviors, leading to a nonhollow structure, microscaled hollow structure or microporous structure formation. Through dissipative particle dynamics simulation, it was discovered that the microstructures (nonhollow, hollow or porous) were determined in the emulsion state by the density of the hydroxyl groups located at the end of each polymer arm (Figure 6). Therefore, SS-PLLA can be synthesized to satisfy both nanofiber formation and micropore structure formation requirements in order to self-assemble into a unique nanofibrous spongy microsphere (NF-SMS), which integrated NF and microporous structures into microspheres for the first time (Figure 6). The deduced mechanisms provide crucial guidance on subsequent new polymer synthesis for the simultaneous control of nano- and microstructures of spheres. For example, a recent study synthesized a new polymer, star-shaped PLLA-block-polylysine (SS-PLLA-b-PLYS) which was also capable of forming NF-SMS [151]. The lysine block of the new polymer was incorporated to provide the necessary stabilizing forces for the porous structure formation, while the PLLA block drove nanofiber formation. In this study, NF-SMS were used to deliver human dental pulp stem cells (hDPSCs) to repair dentin tissues [151]. The nanofibers and micropores were shown to enhance the proliferation and odontogenic differentiation of hDPSCs both in vitro and in vivo, compared with NF-MS without pore structure and conventional solid microspheres (S-MS) with neither nanofibers nor pore structure (Figure 7). Great strides have thus been made in developing methods and understanding mechanisms for not only creating microspheres with NF topologies that mimic ECM and promote a variety of desired cell behaviors, but also the ability to fine tune their microscopic morphology via chemistry to customize porosity.

Figure 6. . Dissipative particle dynamics (DPD) simulations and scanning electron microscopy (SEM) observation of 16-arm star-shaped polymers of varying arm lengths and the formation of different emulsions/microspheres.

As the polymer arm decreases its length, the structures undergo a transition from: (top row) nonhollow; to (middle row) hollow; to (bottom row) spongy. Left column: polymer isosurface from DPD simulation, with individual hydroxyl beads on the bottom half of the droplet shown in red. The conformation of a single SS-PLLA is shown in the square box. Middle column: the internal structure of the same droplet in the left from DPD simulation, with glycerol (water phase) in purple. Some hydroxyls (red beads) are removed for clarity. Right column: the SEM micrographs of representative NF microspheres formed at each scenarios, including NF-MS on the top, NF-HMS in the middle and NF-SMS on the bottom.

Adapted with permission from Zhang et al. [150]. © John Wiley and Sons (2015).

Figure 7. . NF-SMS enhanced dentin sialophosphoprotein protein expression of human dental pulp stem cells in vitro.

DSPP immunofluorescence staining of hDPSCs on NF-SMS, NF-MS and S-MS after odontogenic induction for 4 weeks. Blue: nuclei; green: DSPP; and red: F-actin. The images show that DSPP was detected not only on the surface of NF-SMS but also inside its inner pores; DSPP was only detected on the surface of NF-MS; DSPP staining was negative in S-MS group. Scale bars: 10 μm.

DSPP: Dentin sialophosphoprotein; hDPSC: Human dental pulp stem cell.

Adapted with permission from Kuang et al. [151]. © John Wiley and Sons (2015).

Functionalization of nanofibrous hollow microspheres

In the native tissue microenvironment, cells are exposed to both the nanotopographical cues of the ECM and the chemical cues of the soluble and ECM-bound growth factors. These combined physical and chemical signals of the ECM may synergistically affect cell behavior [16,152]. Although synthetic polymers may lack the ability to specifically interact with cells like natural ECM does, surface functionalization is an effective strategy to counter this shortcoming. By introducing desired chemistries at the interface between cells and microspheres, researchers can present cells with environments of bound biological moieties, with which cells can interact to promote proliferation or differentiation. The surface coating of natural polymers or signaling molecules onto synthetic polymer microspheres sounds straightforward, but may compromise the delicate biomimetic physical features at the micro- and nanoscales. Click chemistry has received increasing interest among various conjugation techniques due to its orthogonality, high efficiency and ambient reaction conditions [153]. Thiolene and thiolyne conjugation techniques are particularly useful for cell carriers since they are biocompatible and require no toxic transition metals as catalysts [154,155]. Compared with adhesion proteins such as fibronectin or laminin, cell adhesion peptides [29,31] may likely be advantageous because the short sequences are less susceptible to bioactivity loss due to conformational changes than in large proteins. Moreover, the ligand density on the microspheres can be more precisely tuned to optimize cell–microsphere interactions for tissue regeneration.

We recently developed functional nanofibrous hollow microspheres (FNF-HMS) by synthesizing a biodegradable copolymer, poly(L-lactic acid)-graft-poly(hydroxyethyl methacrylate) [149]. The copolymer composition was adjusted to enable its self-assembly into FNF-HMS. The introduced acrylic groups on the surface of the FNF-HMS enabled bioconjugation between FNF-HMS and peptides through a one-step ‘thiolene’ click reaction, thus achieving the presentation of both biophysical NF architecture and biochemical peptide cues on microspheres for the first time (Figure 8). A TGF-β mimicking peptide or a BMP-2 mimicking peptide was conjugated onto the FNF-HMS to induce the chondrogenic or osteogenic differentiation pathway in rabbit bone marrow-derived MSCs. The nanofibers and the conjugated peptides were believed to exert synergistic effects, leading to enhanced stem cell differentiation and cartilage or bone regeneration in vivo without using growth factors.

Figure 8. . Functionalization of NF-HMS.

(A) A schematic illustration showing the control over the density of functional acrylic groups presented on the surface of FNF-HMS via the manipulation of copolymer compositions (HEMA block percentage). A higher HEMA block percentage (AD value) leads to a higher density of functional groups presented on the surface of FNF-HMS, and thus a higher density of peptides that can be conjugated. (B) FITC spectrum of FNF-HMS before and after conjugation with Cytomodulin peptides, showing the disappearance of peaks corresponding to acrylic groups (circled). (C) Amino acid analysis of Cytomodulin-conjugated FNF-HMS with different conjugation densities.

Adapted with permission from Zhang et al. [149]. © John Wiley and Sons (2014).

Future perspective

The design of cell therapies with maximal efficacy and minimal surgical invasion presents both opportunities and challenges. Various injectable cell microcarriers have been synthesized to achieve these two goals, among which microspheres have been demonstrated to provide injectability, controllable biodegradability and capacity for drug incorporation and delivery. With new discoveries in biology and nanotechnology, scientists are beginning to design novel nanostructured microspheres incorporating various biomimetic and cell-instructive characteristics. Organic/inorganic nanocomposite microspheres, mimicking the composition and structure of mineralized tissues at the nanoscale, enhance osteoconductivity and bone regeneration. Drugs and biomolecules have been assembled and incorporated into the nanodomains of the microspheres and delivered to cells to regulate cell proliferation and differentiation for the targeted tissue regeneration. NF microspheres, mimicking the nanofibrillar structure of the natural ECM, can mediate cell–matrix interactions and cell function. Moreover, self-assembly of NF hollow microspheres and NF spongy microspheres could further increase porosity and reduce the amount of degradation products, as well as accommodate cells both on the outer surface and inside the pore for increased cell–material and cell–cell interactions and subsequent uniform, continuous tissue formation. In addition, protein-derived molecules can be conjugated onto nanostructured microspheres to direct stem cell fate.

The utilization of nanostructured microspheres as cell carriers is still in the early stages of development, but this pioneering work has already shown exciting results and opened up new opportunities for injectable scaffolding in tissue engineering. We expect that in the coming decade, future studies will explore various nanotechnology and engineering techniques to manipulate the structure and composition of injectable cell carriers on the micro- and nanoscales. More research activity will be devoted to utilizing nanostructured injectable carriers to guide stem cell differentiation and target tissue regeneration.

Executive summary.

• Organic/inorganic nanocomposite microspheres can mimic the composition of natural hard tissues and enhance hard tissue regeneration.

• Conventional smooth-surface microspheres can be used as injectable materials for drug encapsulation and cell delivery.

• Biomimetic nanofibrous structure and beneficial microstructures can be built into microspheres and enhance cell–material and cell–cell interactions, leading to improved tissue regeneration outcomes.

• Cell-instructive biochemical signals can be presented on nanofibrous microspheres to exert synergistic biophysical and biochemical signals to control stem cell fate and direct tissue regeneration.