Accepted Manuscript Core-shell designed scaffolds for drug delivery and tissue engineering Roman A. Perez, Hae-Won Kim PII: DOI: Reference:

S1742-7061(15)00123-3 http://dx.doi.org/10.1016/j.actbio.2015.03.013 ACTBIO 3626

To appear in:

Acta Biomaterialia

Received Date: Revised Date: Accepted Date:

28 October 2014 3 March 2015 8 March 2015

Please cite this article as: Perez, R.A., Kim, H-W., Core-shell designed scaffolds for drug delivery and tissue engineering, Acta Biomaterialia (2015), doi: http://dx.doi.org/10.1016/j.actbio.2015.03.013

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Core-shell designed scaffolds for drug delivery and tissue engineering

1,2

Roman A. Perez , Hae-Won Kim

1

1,2,3,*

Institute of Tissue Regeneration Engineering (ITREN), Dankook University, Cheonan 330-714, Republic

of Korea 2

Department of Nanobiomedical Science & BK21 PLUS NBM Global Research Center for Regenerative

Medicine, Dankook University, Cheonan 330-714, Republic of Korea 3

Department of Biomaterials Science, College of Dentistry, Dankook University, Cheonan 330-714,

Republic of Korea

----------------------------*Corresponding author: Prof. H.-W. Kim (e-mail: [email protected]; Tel: +82 41 550 3081)

For: Acta Biomaterialia

Abstract Scaffolds that secure and deliver therapeutic ingredients like signaling molecules and stem cells hold great promise for drug delivery and tissue engineering. Employing a core-shell design for scaffolds provides a promising solution. Some unique methods, such as co-concentric nozzle extrusion, microfluidics generation, and chemical confinement reactions, have been successful in producing core-shelled nano/microfibers and nano/microspheres. Signaling molecules and drugs, spatially allocated to the core and/or shell part, can be delivered in a controllable and sequential manner for optimal therapeutic effects. Stem cells can be loaded within the core part on-demand, safely protected from the environments, which ultimately affords ex vivo culture and in vivo tissue engineering applications. The encapsulated cells experience three-dimensional tissue-mimic microenvironments in which therapeutic molecules are secreted to the surrounding tissues guided by semi-permeable shell properties, consequently acting as reservoirs for the generation of therapeutics. Tuning the material properties of the core and shell, changing the geometrical parameters, and shaping them into proper forms significantly influence the release behaviors of biomolecules and the fate of the cells. This topical issue highlights the immense usefulness of core-shell designs for the therapeutic actions of scaffolds in the delivery of signaling molecules and stem cells for tissue regeneration and disease treatment.

Keywords: Core-shell design; Therapeutic scaffolds; Drug delivery; Cell encapsulation; Tissue engineering; Fibers; Nano/microspheres

1. Introduction Over the past decades, scaffolding materials have been developed to deliver therapeutic molecules and cells, with the goal of repairing and reconstructing diseased and defective tissues. A series of material actions that favor and even stimulate biological processes in the body in terms of orienting protein adsorption, guiding cellular anchorage and migration, and driving progenitor character into specified cellular lineages, have thus been substantially researched [1–5].

Some key solutions on how to secure drugs and protein molecules within the composition and controllably deliver the molecules to the target sites have been sought to overcome the innate power of synthetic scaffolding materials [3,6,7]. Furthermore, how to safely load tissue cells at relevant quantities, particularly preserving potent stem cell characteristics, and effectively transplanting them into the damages that need regenerative actions, have also been the key scaffold-based obstacles in engineering tissues [3,8,9].

Among the accumulative technologies of scaffolding materials for therapeutic molecules and stem cells, the core-shell design has emerged as a promising approach. The cored inner layer with a shelled outer layer, a typical design feature, enables a number of faceted characteristics that have potential for the delivery system and tissue engineering. Typically, the therapeutic molecules and cells can be secured and partitioned within the layered core-shell structure, and can be delivered to target defects. Furthermore, the capacity to load molecules and cells safely, the tunability of material compositions and parameters to proper shapes and sizes, and the secreting actions of the design, are the beneficial faces of the core-shell structure for use as scaffolding systems. Various shapes, including fibers and spheres, are possible by utilizing different techniques; and a wide range of sizes from submicrometers to millimeters allow for the application of the system in diverse fields from the delivery of small therapeutic molecules to the transportation of tissue cells. Furthermore, the fine-integration of the core-shell building blocks enables tissue-level engineering of cells and extracellular matrices (ECMs).

Consequently, core-shell designed systems are considered to have a multi-faceted nature, providing on-demand biomaterial platforms for drug delivery and tissue engineering.

Here, we summarize the core-shell designed scaffolds and materials that find useful applications in drug delivery and tissue engineering. While many different types of scaffolds and nanomaterials have been developed separately, this review is intended to collect for the first time the information available under this theme. It is thus reviewed first to describe general features of core-shell structures, then to identify the production methods and the typical types, and finally, to detail the applications of the core-shells in drug loading and delivery as well as in cell encapsulation and tissue engineering.

2. General feature of core-shell structures The core-shell structure features two discrete parts, with one inner part (‘core’) and the outer part (‘shell’) completely enclosing the inner portion. As they are partitioned in space, each core and shell can perform independent functions, such as incorporating two different molecules. However, they are both interfaced and molecular-permeable, thus molecular interactions between them are possible, and one side can affect the other. The core-shell can be shaped in either a continuous or a discrete manner. The former gives rise to a fibrous shape whilst the latter generates a spherical form. Only when the size (diameter) of the core part is large enough (tens to hundreds of micrometers) does it allow for loading cells. On the other hand, exogenous signaling molecules can be incorporated without regard to the core size and can be loaded in the shell as well. The role of the core is thus considered as either to load therapeutic molecules for delivery or to hold tissue cells and provide them with 3D culture environments. The shell protects the inner biological ingredients, governing the release kinetics of the core-contained molecules and protecting the viable cells. Fig. 1 depicts the general features of the core-shell designs that are useful for cell-encapsulated tissue engineering as well as for the delivery of therapeutic molecules.

3. Methods to prepare core-shell structures

To generate core-shell structures, the general approach is either top-down or bottom-up. The top-down approach involves an apparatus either designed simply with co-concentric nozzles or more powerfully with microfluidic devices. While a continuous nozzling of core and shell components produces fibrous shapes, the periodic discontinuity in the device generates discrete particulate forms. The top-down approach can generate core-shelled fibers and spheres easily with hundreds/tens of micrometers and sometimes even down to a few micrometers. Conversely, the bottom-up approach utilizes chemical reactions in confined conditions to form phase-separated core-shelled particulates, mainly nanoparticles. This is possible either by in situ phase-separation reactions or by the post-shelling approach. This section outlines the top-down and bottom-up approaches to generating core-shell structures.

3.1. Co-concentric nozzle extrusion

Concentric nozzle is an easy and simple approach to obtain two well-defined structures in a coaxial manner. In order to do so, the basic design consists of a central nozzle, usually made of metal to avoid deformation during injection, around which a second nozzle larger in diameter is placed. The design can be achieved through the use of highly sophisticated nozzles or with homemade nozzles. Hydrogel injection and electrospinning based techniques often require these types of nozzles that allow the injection of two different solutions allocated in well-defined compartments. In order to be successful, the injection must take place at the same speed for both solutions. The hydrogel can be formed into microspheres if the flow is disrupted periodically after the injection.

3.2. Microfluidics generation

A more recent approach that uses microfluidics can only be applied to hydrogel based stems. The system by which microfluidics works allows for the use of small size channels in the preparation of perfectly designed and centered structures. In general, emulsions and other systems have the disadvantage of having broad particle size distributions, making the load

difficult to control. Nevertheless, a droplet based system is able to overcome these limitations. The main principle behind microfluidics is the use of micrometer scale channels that create a stream of polymeric precursor solution. This solution is allowed to flow continuously through one of the streams, but is then broken in a controlled manner by interaction with a second flow composed of an immiscible fluid that will act as the continuous phase. By controlling the viscosities, flow rates and dimensions of the channels among others, microspheres are produced which are then able to gel in the formed shape[10–12]. In the case of the core-shell microspheres, the shell is usually made of the gelling agent, whereas the core is able to carry the desired payload, such as cells or drugs, in the desired carrier material [13]. This technology has also allowed for the fabrication of core-shell structures using stimuli responsive materials that are able to release the payload according to external stimuli [14]. The incorporated anticancer drugs within core-shell particles can be released selectively by the surrounding pH change, which is effective for the selective treatment of cancer or chronic wounds [14].

3.3. Chemical confinement reactions

Chemical reactions are mainly applied for the preparation of core-shell nanoparticles. Two main routes have been described to fabricate them. The first route consists of the preparation of core particles which are then surface-modified to allow for coating with the shell material[15–20]. The second process consists of the synthesis of the core particle in situ followed by the synthesis of the shell coating [21]. Both methods consist of a coating to form the shell, although the second method incorporates the use of specific reagents with growth inhibitors, forming the shell after the core reaction is complete. In both cases, it is important to achieve homogenous shells in terms of thickness and composition. Different methods, such as in situ polymerization, microemulsions and sol-gel reactions have been developed for the preparation of core-shell nanoparticles. For a more detailed explanation of the processes, we recommend an extensive and detailed review on the subject [22].

4. Types of core-shell structures Depending on their size and shape, core-shell structures can be categorized into microfibers, nanofibers, micro/nanospheres, and 3D assembled/constructed scaffolds, as illustrated in Fig. 2. While all forms allow for biomolecular loading and delivery, only the cell-compatible sizes enable cell delivery. In particular, the 3D assembly and construction of fibers or spheres through 3D printing technology provides potential opportunities with core-shell units applicable for tissue engineering. The different types of structures, the methods to prepare them, and the types of materials are summarized in Table 1.

4.1. Microfibers

Microfibers with core-shell structures can have the primary purpose of cell delivery through the core section. For this method, hydrogels are among the most promising material choices. Compared to dense materials, hydrogels have excellent capacity to hold cells within internal networks and to preserve their viability. The cells within the core can be protected from the external environments, as a result, the cell death due to pH changes or oxygen tension can be reduced [23,24]. The physico-chemical properties of the hydrogel fibers determine the status of cells, particularly their phenotypes. The composition and physical stiffness of the hydrogels thus needs careful consideration. Another benefit of cell delivering hydrogel fibers is their shapeadaptability to complex defects which enables the delivery of cells into specific wound sites. Therefore, while the core should be cell compatible, the composition of the shell should be mechanically stable. Alginate is thus among the most commonly used polymers for the shell as it can crosslink easily in divalent solution while maintaining the hydrogel shape [25,26]. On the other hand, many cell compatible hydrogels can be used for the core including collagen, hyaluronic acid, and fibrin[26–28].

Another unique approach in core-shell systems is the use of hollowed hydrogel fibers. These hollow conduits can be used as channels that mimic the structure of blood vessels,

allowing for body fluids and cells to pass through. The design principle is the same as the microfiber hydrogels except for the lack of a core solution [29,30].

The core-shell structure also allows for the encapsulation and release of signaling molecules and the release profiles can be controlled through a diffusion mechanism. The two well-distinguished compartments allow molecules to be encapsulated either in the core or in the shell. An additional feature is the possibility of having two different molecules in the core and the shell, respectively, to obtain sequential release patterns. One recent approach used simultaneous extrusion of alginate and tricalcium phosphate to release model proteins in a sustained manner. By changing the amount of tricalcium phosphate incorporated within the core and the shell hydrogels, the release profiles could be effectively tailored [31].

4.2. Nanofibers

Different from microfibers, nanofibers are unable to hold cells within the core portion. Therefore, the core-shell structure is mainly used for incorporating drug molecules. Although the nanofibers cannot encapsulate cells, the nanofibrous network provides excellent support for anchoragedependent cells as a scaffolding matrix. Therefore, drug-loaded nanofibers can be a proper therapeutic tissue engineering matrix.

Nanofiber based systems are usually generated using the electrospinning technique. Electrospinning involves the injection of a material solution through a metal nozzle under DC electrical power. The injected solution is then collected on a conducting substrate with fibrous morphology ranging in sizes from tens of nanometers to several microns. The size and morphology of the fibers can be adjusted by the injection parameters and the collector design. In particular, co-concentric design of a nozzle enables the production of a core-shell structure. The two solutions need to be immiscible to allow phase separation and present two welldistinguished materials. The core-encapsulated drugs will diffuse out through the shell into the surrounding area.

Different material combinations are possible for the core and shell, enabling diversified core-shell structures, as presented in Fig. 3. When hydrophilic natural polymers are used in the shell with hydrophobic synthetic polymers in the core, the shell can favor biological interactions and cellular responses, while the core supports mechanical robustness[32–34]. If the core is selectively removed in an organic solvent, then a hollowed shell can be created. Chitosan hollowed nanofiber with removal of the PEO core is an example of such design [35]. On the other hand, when synthetic polymers that dissolve in organic solvents are placed in the shell with water-soluble natural polymers in the core, the core part can safely deliver therapeutic molecules [36]. Inorganic phases such as silica xerogels can also be shelled on the biopolymer core to create hard shelled flexible nanofibers that may be useful for bone tissue regeneration [37]. Beyond the core-shell double layer, three-layered nanofiber structures have also been generated, which are comprised of a gelatin shell and core, and a PCL phase in the middle [38].

4.3. Spherical forms; microcapsules and nanospheres

Among the different forms of scaffolds and nanomaterials, spheres have been most widely exploited due to their isotropic shape. While the nanospheres (below submicrons) have been used mainly for drug delivery, the microspheres (over tens of micrometers) have also been found to be proper as cell culture matrix. For core-shell nanospheres, the inner part can either be evacuated or filled with different materials.

Many different biopolymers have been used to produce hollowed nanospheres that can deliver cargo drug molecules contained in the hollowed section. The water-in-oil-in-water double emulsion method is one example of a facile method to generate water-soluble drug loading hollowed hydrophobic biopolymer nanospheres such as PLA, PEG and or methacrylated based ones [39–41] (Fig. 4a). Drug release is largely diffusion-controlled, where the drug molecules diffuse out through the shell barriers. The core can also be filled with water-soluble polymers, such as chitosan, alginate or poly(N-isopropylacrylamide) (PNIPAAm), which can also hold the drug molecules [42–44]. Compared to hollow nanospheres, this type of material-filled core-shell nanospheres can sustain the drug molecules for prolonged periods (Fig. 4b).

Apart from biopolymers, inorganic biomaterials have also been developed to be coreshell nanospheres. Calcium phosphates and silica nanospheres are examples of inorganic biomaterials engineered into a hollow form. In particular, with silica nanospheres, the inner core was also filled with other types of inorganic compositions including magnetic nanoparticles, gold nanoparticles, and quantum dots (Fig. 4c). These silica-based core-shell structured nanospheres have unique properties that enable both imaging / diagnosis and drug delivery. For example, the magnetic nanoparticle core / silica shell enables magnetic resonance imaging while the outer silica shell takes up drug molecules due to its high mesoporosity. Likewise, the gold nanoparticle core / silica shell nanospheres allow for photo-imaging befitting from the gold nanoparticles

while

the

mesoporous

silica

can

deliver

therapeutic

molecules.

The

multifunctional properties of core-shell nanospheres, i.e., therapeutic and diagnostic functions, have recently gained significant interest in biomedical fields for disease treatments and stem cell therapies[45–47].

Microspheres with core-shell structures have mainly been applied for the culture and delivery of tissue cells (Fig. 4d). The most popular is the alginate-based core-shell, where the core is generally evacuated to create alginate capsules by means of co-concentric nozzles or microfluidics [48] (Fig. 4e). Electro-dropping has also been used to generate microcapsules. When the core is filled with cell compatible hydrogels, such as collagen and hyaluronic acid in combination with other proteins, then the functions of the core material are not just to support cells in a viable state but also to drive and trigger their secreting molecules, which have therapeutic actions in disease treatments and tissue repair[49–51]. Therefore, the shell is a semi-permeable membrane that controls the diffusion of molecules inside and outside. In this manner, the shell materials are often modified with thin-layered structures implemented by a layer-by-layer assembly through charge-charge interactions that ultimately improves selective diffusion of nutrients and molecules while preserving structural stability for long periods [52][53]. Biocompatible inorganic phases can also be incorporated in the shell to produce hollow microspheres or to produce hydrogel capsules with mechanical robustness while retaining good molecular permeability[54–56] (Fig. 4f).

4.4. 3D assembly and construction

While core-shell structured fibers or spheres can be adequately applied for the culture of stem cells or the delivery of drug molecules for therapeutic actions, 3D constructions with more defined architecture and automated assembly provide promising platforms for tissue engineering. This mainly relies on the bottom-up approach that assembles the fiber or sphere units into 3D macro-structures. It is possible assemble in a 3D form through either direct-writing / direct-deposition of the units or by filling a pre-designed mould. Cells can be co-assembled with the units after encapsulation within the core-part or during the construction process as the outer assembler. A representative example is the robotic dispensing of core-shell fibers [57]. Different material combinations were used to improve the mechanical properties [57]. Hollow fibers can be obtained by selective dissolution of the core or by direct printing of hollow alginate tubes to form scaffolds [58][59] (Fig. 5a). The rationale is that these hollow fibers are useful to direct tissue ingrowth and as drug delivery depots which can then be placed in a 3D architecture to obtain scaffolds with regular macropores as well as artificial vascular like structures. The hollow structure of the fibers does not largely alter the mechanical properties of the scaffolds and also facilitates homogenous populations of cells on the surface of the fibers as well as in the inner lumen sections. Hence, it can be considered for use as a pre-vascular structure for tissue engineering. The microspherical form is an attractive assembly unit that can be constructed with cells into 3D designed tissue-engineered constructs on the macro-scale. These core-shell units are either made of cell/material (cell/collagen beads) or made of different cell types (cell/cell beads) which enhance the potential opportunities for the cells to be generated into 3D complex tissues and can be scaled up in size and quantity [60] (Fig. 5b). After seven days in culture, biopolymer core-shell microspheres that encapsulated growth factors were able to assemble through cellto-cell and cell-to-particle interactions into 3D tissue engineered constructs [61]. In fact, the idea of 3D assembly of spherical cellular constructs dates back to 3D cell spheroids, where cellular spheroids were created to be assembled into tissue-leveled 3D architectures. In this case, the shape of the pre-designed moulds has shown significant influence on the behaviors of the 3D

cell assembly [62,63] (Fig. 5c). Toroid moulds with different sizes and widths allowed for different interactions between cells, dictating their 3D assembly. While a smaller toroid cone separated two cell populations into two different colonies, a bigger sized toroid allowed the interaction and co-assembly of the two cell populations [62]. Different types of cells and cellular combinations can be utilized in the 3D cell assembly [63,64]. Most tissue structures are better formed with pre-vasculature formation, which necessitates the co-assembly of the endothelial cells with progenitor cells like MSCs for bone (Fig. 5d). In a study on utilizing MSCs and endothelial progenitor cells (EPC) [65], the cellular assembly presented different spherical morphologies, and the concentric assembled morphology of both cells showed the highest angiogenic effect [65]. While the study of 3D cell assembly is still in its infancy, its future as a leading technique for ex vivo tissue engineering, particularly at the sophisticated and large-sized tissue-levels, is considered promising.

5. Applications for drug loading and delivery

5.1. Single drug delivery

Core-shell designs are specifically needed for the delivery of drugs and molecules that have different functionality or that should be secured or protected from the processing conditions. For example, water-soluble hydrophilic drugs can be encapsulated in the hydrophilic core material and shelled with hydrophobic material that is soluble in organic solvents. Similarly, the opposite is also possible, i.e., hydrophobic drugs in the hydrophobic core polymer with a hydrophilic shell layer. While individual fibers and spheres without core-shell structures have been commonly used for single drug delivery purposes, the shell layer can also control and slow down the release of the loaded drugs [66].

Nanofiber designs have long been combined with core-shell structures aiming to deliver therapeutic molecules. Changing parameters such as the composition of the core and the shell, the chemical properties, as well as the injection speeds, showed profound influence on the release profiles of the delivered molecules [67][68]. A hydrophilic core with a hydrophobic shell

is a proper combination for the delivery of proteins and hydrophilic drugs. Many growth factors have been used in the core-shell nanofiber delivery system for the repair and regeneration of tissues including bone, muscle and nerve. Polycaprolactone (PCL)-based electrospinning fibers effectively encapsulated a model protein in the core [69]. The shell size could be increased by increasing the feeding speed of the PCL solution without significantly affecting the fiber size and the increased shell reduced the release rate of the protein. Similarly, polyethylene glycol (PEG) was used to encapsulate BMP2 within the core of a PCL nanofibrous structure, achieving a zero order release for over 24 days. Furthermore, osteogenic genes were expressed better by the cells in vitro as well as in vivo on the BMP2-delivered nanofibers [70]. When PEG was blended with PCL in the shell, pores could be generated, demonstrating pore-controllable BMP2 release [71] (as illustrated in Fig. 6a). The addition of PEG to the shell could control the protein release between 1 week and 1 month [72]. In a similar manner, the incorporation of VEGF into PCL and PLGA nanofibers showed excellent activity on the performance of endothelial cells [73][74]. The core-shell gelatin-based nanofiber was tailored with heparin immobilization to control growth factor release behaviors [75].

PCL shell with dextran core also demonstrated controllable delivery of protein molecules [72]. The PLCL and dextran core-shell nanofibers encapsulating platelet derived growth factor (PDGF) showed that the inner flow rate could affect the amount of growth factor loaded. The PDGF release was able to enhance the attachment and cellular activities of smooth muscle cells [76]. The PLCL-dextran combination was also useful for the loading and delivery of TGF-β1 [77]. Nerve growth factor (NGF) was also incorporated within the core-shell PLLA conduits and implanted into a 10 mm gap of rat sciatic nerve, revealing comparable results to autograft control [78]. Likewise, the release of NGF from PLGA showed an enhanced regeneration in vivo, which was potentiated when the nanofibers presented an aligned morphology [79].

Similar to the fibers, sphere forms have also been generated to deliver drugs and proteins. Polymeric core-shell microspheres based on PHBV and PLGA showed highly sustained release of hepatocyte growth factor over 40 days [80]. PLGA and PLLA core shell microspheres also allowed sustained release of bupivacaine in vivo during a 2 week

implantation in goat knee [81]. In a slightly different approach, a hydrogel based core sheathed with a polyester shell allowed for improved loading of hydrophilic drugs as well as sustained release [82]. In fact, the presence of the core hydrogel allowed for sustained release, although the thickness of the shell as well as the material used have also revealed significant effects on the cargo release [83]. Sometimes, nanoparticles were incorporated into core-shell microparticles for more sustained release [61]. Self-assembling peptides were also used to form hydrogel microspheres to load and release antibodies [84] (as illustrated in Fig. 6b). The concentric spheres were formed using ac-(RADA)4-CONH2 and (KLDL)3-CONH2 as core and shell structures, respectively, and the release was sustained over 3 months, with the biological activity of the encapsulated antibody remaining intact [84].

Compared to microspheres, that can be used for the delivery of drugs and proteins to be released outside of cells, nanoparticles are mainly intended to be delivered into cells, and thus the drug delivery purpose can be more potentially benefited by the nano-sized particles. For this reason, the delivered therapeutic molecules are designed to function inside cells, and include genes, proteins and drugs that are more relevant in direct interactions with intracellular compartments. Core-shell designed nanospheres thus function most effectively in loading and safely delivering through cellular permeating processes while maintaining targeted and safe intracellular actions [43][85]. Compared to the dense nanospheres, which often utilize surfaceloading of drug molecules, core-shell systems can hold molecules in the core (‘nanocontainers’), which are hollow-spaced or matrix-bound, at much higher quantities and with greater safety.

For the hollow nanoshells, the encapsulation of drug molecules is possible in the course of or after shell formation, and drug release can be either from the rupture of the shell composition, or diffusion through the shell [86]. In particular, the rupture mechanism is often designed to be stimuli-dependent, in such a way that it is responsive to temperature, pH, or enzyme [87][88] (as presented in Fig. 6c). For example, it was previously shown that mesoporous silica nanoparticles coated with pNIPAAm presented a temperature dependent release profile, with significantly higher release at 37ºC than at room temperature [89][90]. In a more biological approach, DNA was used as a coating for the nanoparticles, showing that when the temperature was increased and the DNA was denatured, the cargo within the nanoparticle

was released [91]. Similarly, certain polymers (e.g. polyethylene glycol diacrylate) that are protease-sensitive can be coated on the surface of nanoparticles, ensuring that the incorporated cargo was released only when placed in contact with the protease present in cells in vitro or in vivo [92]. On the other hand, when a hydrophilic polymer was used for the core and layered with a thin hydrophobic shell, the inner matrix held drug molecules more selectively and sometimes more tightly, releasing them in a more controllable and sustained manner [93].

Core-shell nanoparticles can also be designed to exert multifunctional activities through specific actions, such as imaging and sensing, of the core and/or the shell [95-98]. Some representative core-shell nanoparticle designs have recently been developed with the purpose of enabling multifunctionality including diagnostics and therapeutics. One design is the use of MSN as the core and a polymer shell with stimuli-responsiveness, such as temperature and magnetism. MSN can be coated with pNIPAAm, showing a temperature dependent release of DOX [95][89,92]. Likewise, a zwitterionic sulfobetaine copolymer was coated on the MSN, causing the release to be temperature dependent [96]. Higher temperatures triggered volume contraction of the polymer shells and accelerated the diffusion which in turn released the molecules inside faster [96]. Furthermore, when an external magnetic field was applied to coreshell nanoparticles designed with a hydrophobic transformable core and an ultra-thin shell, the increase in temperature induced shrinkage of the core and rapid drug release [97]. Another representative design includes a mesoporous silica shell with imaging-available nanoparticles in the core (magnetic nanoparticles, gold nanoparticles and quantum dots)[98–100]. For example, superparamagnetic nanoparticles coated with thin mesoporous silica layers enabled magnetic resonance imaging while the mesoporous silica effectively loaded and delivered drug molecules [99][101]. Similarly, gold-nanoparticles shelled with mesoporous silica facilitated CT-scanning through photo-thermal activation of gold while photo-thermal therapy using gold was also enabled[102–105] (as depicted in Fig. 6d).

While the core-shell designs described above have mainly been used for the delivery of single drug molecules, their utility can extend the delivery of multiple drugs with different and synergistic therapeutic actions, which will be discussed in the following section.

5.2. Multiple drug delivery

Although individual drugs can have therapeutic actions when delivered at proper doses and for the proper time period, multiple actions of different types of drugs delivered at the same time or in a sequential manner can potentiate and synergize the therapeutic capacity [106][107][5][3]. Consequently, many recent studies have focused on multiple drug delivery systems. In the tissue repair and healing process, a cascade of multiple signaling molecules is generally engaged in combination with other processes in a time-dependent manner[108]. For example, rapid homing of progenitor cells accompanied by tissue regeneration, angiogenesis followed by osteogenesis, and initial anti-inflammatory action with later tissue recovery, are the most common events that occur in the repair and regenerative processes involved in tissueimplantable materials [108][109]. Therefore, delivery systems that can mimic better the in vivo conditions, i.e. can secrete therapeutic molecules in a time-dependent manner, are considered to be more effective for tissue performance. The core-shell structure has merits for this purpose. Different types of molecules can be incorporated within the core and shell structures, respectively, where the drug in the shell releases more rapidly than the molecule encapsulated in the core. The shell that encloses the core can effectively sequester the therapeutic molecules inside and can substantially retard molecular diffusion. In this manner, the materials chemistry and properties, such as degradation rate and molecular interactions, are of special importance. Some recent studies have demonstrated modeled experiments of core-shell designs in the controllable delivery of different molecules. A hydrogel core-shell fiber based on alginate was combined with bioactive ceramic nanopowders targeted for bone, and a model protein cytochrome C was encapsulated either in the core or in the shell. Results showed that the protein incorporated in the core presented much slower release compared to that in the shell and that the release could be tailored depending on the amount of nanopowders incorporated [31]. Using electrospun nanofibers, a similar work was also conducted, where a more sustained release of protein molecules was observed when incorporated in the core [110]. Another interesting study was also performed, where the VEGF was loaded in the inner part of the membrane while PDGF was loaded in the outer part [111]. The goal was to initially enhance proliferation of vascular endothelial cells on

the inner compartment followed by the proliferation of vascular smooth muscle cells in the exterior compartment. Results showed that this system enabled proper proliferation of each part of the cell as well as the development of endothelial cells on the lumen and smooth muscle cells in the exterior of the rabbit carotid artery [111] (as illustrated in Fig. 7). This work signifies the impact of spatial localized design of therapeutic molecules to deliver successive scaffold-based engineering of tissues, and the possibility of state-of-the-art delivery design for tissue engineering. In fact, the core-shell fibrous structure has been well demonstrated to be effective in delivering multiple drugs simultaneously. Electrospun fibers of gelatin-combined copolymers with either PLA or PCL in the core and shell, respectively, are an example. Several epidermal growth factors were incorporated into the core of the core-shell structure. Results showed that the core-shell structure allowed a more sustained release with no initial burst, whereas the non core-shell structured fibers presented an initial burst of as high as 44.9% within the first 15 days [112]. A similar result was found in the delivery of BMP2 and dexamethasone, which were both incorporated within the core and exhibited a highly sustained release of both factors in osteogenesis sufficient for bone regeneration [113]. The shell is thus considered a simple and effective diffusion barrier between the cored therapeutic molecules, prolonging the release profile and improving the biological functions [114]. Microspherical capsules have also been used as delivery vectors though core-shell designs [4]. A recent work described the use of PLGA at different molecular weights to control the release profile from the core while alginate was used in the shell. The microcapsule successfully encapsulated PDGF within the core and VEGF in the shell, and demonstrated a significantly faster release of VEGF compared to PDGF [52]. A co-axial electro-dropping method was also used to obtain microcapsules comprised of PLGA-core / alginate-shell, which incorporated BMP2 and dexamethasone in the core portion. Results showed that the initial burst decreased considerably when the biomolecules were loaded into the core. Furthermore, when MSCs were cultured for up to 4 weeks, a series of osteogenic markers were considerably enhanced by the co-delivered osteogenic factors [115] (as illustrated in Fig. 8).

As demonstrated, core-shell designs enabled the fine-tuned delivery of bioactive signaling molecules in combination or sequentially. The delivered molecules from the designed scaffolds have demonstrated the ability to exert therapeutic actions that stimulate cellular responses and generate neo-tissue formation. In the following section, another area of coreshell designs in cell encapsulation and delivery is detailed.

6. Applications for cell encapsulation and delivery Here, the purpose for the use of core-shell biomaterials in the encapsulation and delivery of cells is two-fold: one is to utilize the encapsulated cells for the secretion of specific signaling molecules that can further stimulate other cells to cure diseases and remedy defects, and the other is that the encapsulated cells and the scaffold constructs are considered as tissue engineering platforms implanted as a whole in vivo with the goal of releasing the cells in the defect sites to reconstruct and regenerate the tissues.

6.1. Therapeutic actions of cells encapsulated

The microencapsulation of cells has been widely applied as a facile tool for transplanting cells, including stem cells [116]. Microcapsules allow for the diffusion of nutrients and signaling factors inwards to the cells, releasing at the same time signaling molecules outwards to the surroundings. The physical properties of microcapsules, including high surface area to volume ratio, resistance to mechanical stress, short diffusion lengths, and their ability to be processed, positions them as a proper choice for cell delivery vehicles. The encapsulation needs to be precise and accurate, since the exposure of a single cell often creates an immune response that causes complete rejection of the transplant [117]. For this reason, the properties and choice of shell materials should be carefully considered to guarantee that the microcapsule structure does not collapse during implantation for a specific period of time [118].

In fact, the most commonly transplanted cells using microencapsulation are the pancreatic islets for diabetes mellitus treatment. Changes in blood glucose levels diffuse from the blood vessels into the core-shell capsules, allowing for the control of glucose metabolism. Alginate is the most commonly used material [119][120]. Many divalent cations including CaCl2 and BaCl2 can produce crosslinking of alginate and the degree of crosslinking varies depending on the cation type [120]. Even thin alginate shells (10 µm) preserve the islets’ functionality and viability, without any lag in insulin secretion from the capsules [120]. The thickness was reduced (1.5 µm) even more through the use of a synthetic polymer, but the cell viability was reduced due to the cytotoxic nature of the polymer [51]. Agarose gel was also used for cell encapsulation and significantly improved the cell viability [49]. Multiple step processes are often used for shell formation, which is not optimal for cell survival and thus the shell part is often unable to withstand in vivo degradation [50]. For this reason, a more sophisticated method was recently used where two fluid co-axial electro-jetting was used in a single step that concentrates cells in the core with a stable shell layer and better control of the capsule size and morphology [121] (as illustrated in Fig. 9).

6.2. Tissue engineering

As described above, the microencapsulation of cells in core-shell designs was first aimed at treating diseases, where the therapeutic molecules secreted from the encapsulated cells were effective in treating incurable diseases like diabetes. The cells encapsulated are generally in the form of aggregates, with the inner core structure having a minimal role, presenting the shell as protector and molecular transfer vehicle.

Evolving from this microencapsulation concept, some recent approaches have emerged to deliver cells for tissue engineering purposes, where the selection of appropriate core materials is a key aspect. The properties of the core materials, such as composition or stiffness, can dictate the fate of the encapsulated cells, ultimately directing their proliferation or differentiation behaviors to be favorable for tissue regeneration. A major focused cell source for this is stem cells with either pluripotency or multipotency. While embryonic stem cells (ESCs)

are generally in the form of aggregates with a similar feature as the embryo body, multipotent cells are well dispersed in the core matrix. In fact, maintenance of the pluripotency of ESCs while retaining a high survival rate and proliferative potential has been a challenging issue in current ESC research [122][123]. For this purpose, some recent works encapsulated ESCs into microcapsules made of alginate [124][125]. A non-planar microfluidic system was introduced for ESC microencapsulation. Results showed that the microencapsulated ESCs were able to maintain their pluripotency to a higher extent when compared to those conventionally cultured on 2D culture dishes, signifying the 3D fluid-like culture conditions were effective in preserving the ESCs’ pluripotency, overcoming the limitation of conventional 2D cultures, and providing possible technological approaches for better use of ESCs (as illustrated in Fig. 10).

Recent works have focused on control of the properties of the core materials. Cells that are encapsulated precisely sense the internal physical and chemical cues in the core material, such as the chemical groups and mechanical stiffness. In particular, the stiffness of core materials has been shown to be a significant factor in differentiation of the encapsulated cells [3]. An elegant approach was recently reported where different ECMs were introduced into the core [25]. A microfluidic device involving a double coaxial laminar flow was used to obtain meter-long microfibers encapsulating different types of cells surrounded by an alginate shell [25]. The different proteins used were pepsin-solubilized collagen, acid-solubilized collagen and fibrin, attaining elastic modulus of 6.3, 154 and 730 Pa, respectively. When cultured in coreshell devices tuned differently, the cells were able to produce a tubular structure composed mainly of cells after digestion of the alginate shell. With this fascinating technology and the combination of different proteins and cells, the authors were able to obtain fiber like structures with very similar tissue functionality [25]. The results showed that the ECM-like fiber formation was only successful when the cells, including fibroblasts (3T3), endothelial cells and myocytes, were cultured in relatively stiff substrates, but not in the soft substrate. The cells were unable to properly adhere to and proliferate on the low stiffness substrate and consequently could not form ECM molecules. On the other hand, stiff substrates were favorable for cells to secrete their own ECM molecules, allowing for the formation of well-organized ECM fibrous structures (as illustrated in Fig. 11). Moreover, core-shell structured microcapsules were prepared with a microfluidic device through the self-assembly of peptides of opposite charges [126]. The self-

assembled fibrous microcapsules were found to be stable in aqueous solutions and their permeability was dependent on the capsule composition. Furthermore, the human dermal fibroblasts encapsulated remained viable within the microcapsules and their morphology was shown to be influenced by the matrix density. While the fibroblasts presented a spherical morphology with poorly organized F-actins at high peptide concentrations, the cells tended to have a spindle shape with lots of protrusions at low peptide concentrations, indicating that the cell morphological phenotypes could be controlled by the density of the core materials.

Not only the core material, but the shell was also varied with different compositions. An interesting work utilized silica sol-gel material for the shell portion [127]. A microfluidic-devised core-shell structure was generated with methacrylated gelatin in the core combined with silica in the shell. The role of the silica shell implemented by a sol-gel process was to protect the cells from oxidative stress or mechanical stress that might be caused by the injection. The cardiac muscle cells encapsulated within the core were able to actively populate on the biocompatible and biodegradable microgel surface and then to proliferate over time [127] (as illustrated in Fig. 12). An interesting study that utilized the

Another important consideration of cell delivery core-shell designs is mass transfer and cell survival in the structure. A recent study developed alginate core-shell structure by plotting them into 3D constructs, and showed that the cells close to the border of the shell material were the only ones alive, whereas those placed at the central core were unable to survive [128]. Similarly, a recently designed fiber made of collagen core / alginate shell also demonstrated that rat MSCs encapsulated in the collagen core tended to form active cellular processes at the core/shell border areas to achieve better nutrients and oxygen supplies from the outer environments, which consequently became functional in forming neo-bone structure in vivo [26] (as illustrated in Fig. 13). Therefore, the formation of capillary within the core is considered to be an appropriate approach to follow in future studies.

Thus far, the core-shell designed fiber and spherical constructs have been demonstrated to be effective depots for cellular survival and delivery into target tissues, and the properties of the stem cells to be delivered within the core largely depended on the physical and chemical parameters of the core and shell. Therefore, proper tailoring of the materials and the

combination of signaling molecules within the core-shell compartments may improve the quality control of stem cells and functional outcomes in tissue repair and regeneration.

7. Concluding remarks The use of core-shell systems opens a new door to the field of drug delivery and tissue engineering. The design of biomaterials that are able to release signaling molecules in a controllable manner and to encapsulate and deliver cells effectively significantly improves the regenerative capacity of scaffolds. Different designs are available for core-shell structures, such as nano/microfiber matrices, nano/microspherical particles, bottom-up assemblies and 3Dconstruct tissue mimics. Three main designs are possible, i.e., use of core-shell nozzles, microfluidic systems, and coatings through chemical reactions. Fine control over the processing parameters enables the core-shell designs to be specifically applicable for drug delivery and cell delivery vehicles. Therapeutics delivery by core-shell structures allows safe loading of drug and protein molecules in large quantities. Furthermore, not only one drug, but multiple drugs can be delivered in a sustained and even time-sequenced manner when the drug molecules are loaded in the core and shell structures, selectively. Fibrous delivery systems comprise of various combinations of materials, e.g., synthetic polymer core/natural polymer shell and polymer core/inorganic shell, and the multiple layered structures have been exploited to attract tissue cells or enable sustained drug release. Furthermore, nanospheres with core-shell structures including hollowed, filled or hybridized designs, have been used as drug delivery nanovehicles in cellular compartments. On the other hand, cell delivery with core-shell designs was made possible by the encapsulation of cells within their structure, protecting the cells from external biological environments. Two different approaches have been introduced for cell encapsulation. In one approach, the cells, genetically modified or from allogeneic or xenogeneic origins, can be encapsulated to secrete signaling molecules to induce changes in the environment, such as for the treatment of Diabetes. In the other approach, the cells can be incorporated into the core structure to be developed as tissue-engineered constructs and consequently made implantable

into defective tissue. Their fate can be controlled by adjusting the parameters of the core and shell structures. Present knowledge on biomaterial properties, including chemical groups and physical rigidity, need to be applied in fine tuning of the core and shell materials in order to regulate the fate of stem cells that are to be encapsulated and cultured in the core materials. While core-shell designs have mainly been realized in spherical and fibrous structures which ultimately serve as scaffolding matrices for drug delivery and tissue engineering, they can be formulated into more sophisticated shapes. For instance, recent effort has been taken to fabricate 3D printed scaffolds with core-shell morphology either with or without living cells. The assembly of cellular constructs of core-shell ingredients into large 3D structures has also opened a new pathway to create large scale tissue mimic 3D structures, where cellular and material components mutually interact to enable vasculature and tissue formation. Recent studies have utilized core-shell designs for many applications. Some representative examples include dual growth factor delivering blood vessels, multiple drug releasing hydrogel microcapsules, alginate-based capsules that maintain pluripotency of the stem cells within, ECM-engineered tubular structures for specific tissue differentiation, hybrid cell-encapsulating spheres that relieve oxidative and mechanical stresses, and cell-vascularized microtubes for bone tissue engineering. Beyond these, there still remain more exciting fields in regenerative medicine that can be attained using core-shell designs. In particular, fine tuning of the delivery profile for therapeutic molecules, on-demand dictation of stem cell fate, and ex vivo creation of 3D tissue mimic structures require further development of the technological processes and materials parameters of core-shell designs. While the potential of core-shell systems for the drug delivery and cell encapsulation has been implicated, there are some technological issues that need to be solved and advanced. Although many biomaterial compositions have been developed to encapsulate cells as the core and shell parts, the properties of those materials still need to be improved and optimized for the maintenance and dictation of proper cellular behaviors as well as for the stable isolation of cells with sufficient exchange of ingredients. Furthermore, the core-shell devices should be advanced in much simpler and cheaper ways that ultimately enables practical uses in tissue engineering and drug delivery. Still, most studies of the core-shell biomaterials have focused on the in vitro cellular performance; therefore, future researches need to be directed toward relevant pre-

clinical studies that can ultimately realize the core-shell technologies for practical applications in clinical settings.

Acknowledgements This work was supported by Korea Health Technology R&D Project (HI14C0522) through the Korea Health Industry Development Institute (KHIDI), and Priority Research Centers Program (grant#: 2009-0093829) through the National Research Foundation (NRF), Republic of Korea.

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Figure captions

Fig. 1. Illustration of the general features of core-shell designs that are useful for cell encapsulated tissue engineering (a) as well as for the delivery of therapeutic molecules (b). Fig. 2. Illustration of core-shell structures categorized into (a) microfibers, (b) nanofibers, (c) micro/nanospheres, and (d) 3D assembled/constructed scaffolds depending on their size and shape. Fig. 3. Illustrations of nanofibers with different material combinations for the core and shell. (a) Hydrophilic natural polymer shell with hydrophobic synthetic polymer core favors biological interactions and cellular responses, where the core part can be selectively removed to create hollow shell nanofiber. (b) Synthetic polymer shell with natural polymer core effective for delivery of therapeutic molecules such as growth factors. (c) Inorganic phase in the shell placed on polymer core generates hard and stiff surfaces effective for bone tissue. (d) Multiple layered core-shell structure also possible. Fig. 4. Core-shell based micro/nanospheres: (a) Hollowed nanospheres produced by water-inoil-in-water double emulsion method. (b) Core-shell nanospheres filled with material for sustained delivery. (c) Nanospheres hybridized with nanoparticles for multifunctionality such as imaging. (d) Microcapsules to deliver cells. (e) Alginate capsules produced by microfluidics are representatively shown (with permission from ref [48]) (f) Inorganic phase shelled cell-capsules (exampled from ref. [54] with permission). Fig. 5. 3D assembled or constructed core-shells. (a) Robotic-dispensed core-shell hollowed 3D scaffold (with permission from ref [58]) (b) 3D assembly of core-shell spherical units made of either cell-collagen beads or cell-cell beads, to create large 3D tissues (with permission from ref [60]). (c) Pre-designed mould used to create the 3D cell assembly influences cell behaviors (with permission from ref [62]). (d) Pre-vasculature formation enabled by co-assembly of the endothelial cells (green) with MSCs (red) (with permission from ref [65]).

Fig. 6. Drug delivery applications with core-shell structures. (a) PEG blending with PCL in the shell creates pores showing pore-controllable BMP2 release (with permission from ref [71]). (b) Self-assembling peptides form hydrogel microspheres to load and release antibodies (with permission from ref [84]). (c) Core-shell nanospheres developed for stimuli-responsive DDS. Rupture mechanisms including temperature-, pH-, or enzyme-responsiveness, depend on the shell material. (d) Core-shell nanospheres hybridized with other functional nanoparticles enable imaging, such as CT-scanning for gold/silica nanoparticles. CT images of mouse liver (red arrows), spleen (blue arrows) and kidney (pink arrows) before and after injection of Au/SiO particle colloid solution (with permission from ref [102,105]). Fig. 7. Core-shell hollowed nanofibers designed for delivery of VEGF and PDGF, to enhance initially proliferation of vascular endothelial cells on the inner compartment followed by the proliferation of vascular smooth muscle cells in the exterior compartment. Results showing proper proliferation of each part of the cells as well as the development of endothelial cells on the lumen and smooth muscle cells in the exterior of the rabbit carotid artery (with permission from ref [111]). Fig. 8. PLGA-core / alginate-shell microspherical capsules to deliver BMP2 (core) and dexamethasone (shell). Initial burst decreased considerably when biomolecules were loaded into the core. Furthermore, a series of osteogenic markers were considerably enhanced by the co-delivered osteogenic factors at 4 weeks of MSCs culture (with permission from ref [115]). Fig. 9. Cell delivery core-shell designed using two fluid co-axial electro-jetting enabled in a single step that concentrates cells in the core with a stable shell layer (with permission from ref [121]). Fig. 10. (a) Utility of 3D culture in core-shell hydrogel microcapsules of ESCs to maintain pluripotency, which is often limited in 2D culture in plastic dishes. (b) ESCs cultured in 3D gel (with permission from ref [124]). Fig. 11. Hydrogel microfibers of ECM protein core with alginate shell that carry tissue cells produced by microfluidic device. Stiffness modulated differently (6.3, 154 and 730 Pa). While

cells in soft matrix were unable to form ECM fiber formation, those cultured in relatively stiffer substrates formed ECM fibers (with permission from ref [25]). Fig. 12. Microspherical cell carriers of methacrylated gelatin core with inorganic silica shell by sol-gel process. Silica shell protects the encapsulated cells from oxidative stress, mechanical stress and immune responses that might be caused by the injection. The cardiac muscle cells encapsulated within the core were able to populate actively on the biocompatible and biodegradable microgel surface and then to proliferate over time (with permission from ref [127]). Fig. 13. MSCs cultured in the collagen core-alginate shell fiber cell delivery system form highly networked cellular constructs in vitro and engage in proper osteogenesis (with permission from ref [26]).

Table 1. Summary of various designs of core-shell structured materials developed for the delivery of drugs and cells.

Materials

Type

Micr ofibe r

Nan ofibe r

Fabrication method

Core

Shell

Application

Ref ere nce

Co-concentric extrusion/Fluidics

Collagen, Fibrin

alginate

Cell encapsulation - Fibroblast, myocyte, endothelial cells, nerve cells, epithelial cells

25

Co-concentric extrusion/Fluidics

Collagen, Fibrin

Alginate

Cell encapsulation - Diabetes mellitus treatment

25

Co-concentric extrusion

Collagen

Alginate

Cell encapsulation - Bone tissue regeneration

26

Co-concentric extrusion

Hollow

Alginate

Cell encapsulation - Hollow fibers Cartilage regeneration

29, 30

Co-concentric extrusion

Dextran

Alginate

Cell encapsulation - Islets of Langerhans - Diabetes mellitus treatment

124

Co-concentric extrusion

Alginate+Tri calcium phosphate

Alginate

Dual Drug Delivery - Bone Regeneration

31

Co-axial electrospinning

Poly Vinyl Alcohol

gelatin

Tissue regeneration - Enhanced mechanical properties and cell activity compared to PVA fibers

32

Co-axial electrospinning

Poly (Glycerol sebacate)

gelatin

Myocardium restoration - Enhanced mechanical properties and cardiogenic differentiation

33

Co-axial electrospinning

Poly (Ethylene Oxide)

Chitosan

Purification of blood in hemodialysis and wound dressings

35

Co-axial electrospinning

Hollow

Chitosan

Purification of blood in hemodialysis and wound dressings

35

Co-axial electrospinning

PLGA

Collagen

Bone Tissue Engineering - Dual Drug delivery vehicle (fibronectin/Cadherin 11)

36

Co-axial electrospinning

Poly (Vinyl Alcohol)

PCL, PLLA, PLGA

Drug Delivery Vehicle Optimization of release rates

67

Co-axial electrospinning

PEG

PCL

Controlled release of proteins and drugs for tissue engineering

69, 70

Co-axial electrospinning

PEO

PCL-PEG

Bone Tissue Engineering Sustained drug release

71

Co-axial electrospinning

Dextran

PCL, PLA

Controlled release of proteins and drugs for tissue engineering

72, 74, 76

Micr ophe res

Co-axial electrospinning

pHMGCL, PVPD

PCL

Tissue engineering - Sustained release of growth factors

75, 77

Co-axial electrospinning

Gelatin

PLLCL

Skin regeneration - Dual Drug Delivery system

116

Co-axial electrospinning

PLLCL

Collagen

Bone Tissue engineering - Dual drug delivery vehicle

117

Triaxial Electrospinning

PCL (2nd Core)/Gelati n (1st Core)

Gelatin

Tissue Engineering scaffolds Multiple Drug Delivery vehicle

37

Electrospinning coating

Poly (Caprolacton e)

Collagen, Gelatin

Tissue Engineering - Functional biomimetic nanofiber scaffold

34, 75

Emulsion Electrospinning

Methyl Cellulose

PDLLA

Tissue Engineering Scaffolds Sustained control of encapsulated drugs

68

Droplet coating

Alginate

Calcium silicate

Bone Tissue Engineering Enhanced mineralization and protein delivery control

47

Co-axial Electrodropping

PLGA

Alginate

Angiogenic stimulator - Dual Drug delivery (VEGF and PDGF)

51

Co-axial Electrodropping

PLGA

Alginate

Bone Tissue engineering - Dual drug delivery (BMP2 and dexamethasone)

119

Co-axial electrojetting

Matrigel

Alginate

Islets microencapsulation for type 1 diabetes treatment

125

Layer by Layer coating

PLGA

Layer-by-layer (chitosanhyaluronic acid)

Soft tissue regeneration

52

Biomimetic approach

Gelatin

Calcium phosphate

Bone tissue regeneration drug delivery system

55

Concurrent ionotropic gelation and solvent extraction

PLGA, PLLA

Alginate

Drug delivery - improved loading and release of water soluble drugs

82

W/O/W emulsion solvent evaporation

PLGA

PHBV

Liver tissue engineering - Drug delivery system

80

W/O/W emulsion solvent evaporation

PLGA

PLLA

Postoperative pain release after knee surgery - Delivery of bupivacaine

81

Microfluidics

PLGA

Alginate

Drug delivery - narrow size particles to allows controlled drug delivery

83

Nan osph eres

Microfluidics

Embryonic stem cells

Alginate

3D architecture pre-hatching mimicking environment for embryonic stem cells

128 , 129

Microfluidics

Gelatin

Silica

Tissue engineering - Cell protected environment combined with ceramic phase

132

Chemical confinement reactions

CuInS2

Zn1−xMnxS

Magnetic resonance and fluorescence imaging of cancer cells

44

Chemical confinement reactions

CdSe

Zn1−xMnxS

Dual mode optical and magnetic resonance imaging

45

Chemical confinement reactions

Silica

Chitosan

Applications in ultrasound induced imaging

87

Chemical confinement reactions

Fe3O4, MnFe2O4

Silica

On demand drug delivery and theranostic agent

88, 105

Chemical confinement reactions

PEG-Fe

Fe3O4

Tumor targeting, photothermal therapy and imaging

97

Chemical confinement reactions

Gold nanorods

Silica

Imaging and photothermal therapy of cancer

Nanoparticle capping

Paraffin

Silica

On demand drug delivery

Nanoparticle coating

Fe3O4

Dextran/acryla myde

Magnetic drug targeting delivery and magnetic resonance imaging

Coaxial electrospraying

PVP, PCL

PLGA

Dual drug delivery - Tumor chemotherapy applications

94

W/O emulsion + sol-gel procedure

MnO

SiO2

Dual mode optical and magnetic resonance imaging

46

Selective core dissolution

PELA

Chitosan

Nanocarrier drug delivery system in vitro cell growth inhibition

39

Pickering emulsion polymerization

PNIPAAm

Silica

Cancer treatment - Controlled drug release inducing prostate cancer cell death

41

Reverse microemulsion

Chitosan

Alginate

Gene delivery vehicle - Similar results to those of polyethyleneimine (gold standard)

42

EmulsionPrecipitation

PLGA

Casein

Sequential dual drug release Nanolypharmaceutics platform

43

106 , 107 91

102

3D asse mbly

Core shell nozzle + 3D printing

Hollow

Alginate, PEOT/PBT

Tissue Engineering - Can act as scaffolds for vascular systems

57, 58

Core shell nozzle + 3D printing

Alginate (6%)

Alginate (3.5%)

Tissue engineering scaffold with 3D structure and cell encapsulation

133

Figure1 new

Figure 2

Figure 3

Figure 4

Figure 5

Figure 6

Figure 7

Figure 8

Figure 9

Figure 10

Figure 11

Figure 12

Figure 13