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Cell Encapsulation

  • Abdul Waheed
  • Mohammad Abu Jafar MazumderEmail author
  • Amir Al-Ahmed
  • Partha Roy
  • Nisar Ullah
Living reference work entry
Part of the Polymers and Polymeric Composites: A Reference Series book series (POPOC)

Abstract

The task of developing novel techniques for curing human illnesses is really a complex and tough challenge. This chapter gives a comprehensive discussion of various materials and techniques used in cell encapsulation. Cell encapsulation is a technique whereby living cells are entrapped into a selectively permeable polymeric materials (membranes/beads) making them a potential tool for the treatment of various human illnesses such as Parkinson’s disease, hemophilia, lysosomal storage disease (LSD), cancer and diabetes. The encapsulated cells become immune, i.e., the immune system of the host could not recognize them; therefore, it does not develop any potential immune response against encapsulated cells. Overall, this chapter reviews wide range of techniques that could potentially use in cell encapsulation and discuss how the capsule properties are related to the performance of the cell to treat various diseases. Furthermore, the use of different materials and their impact on the properties and performance in cell encapsulation are also discussed in detail.

Keywords

Cell encapsulation Polymer therapeutics Immune-isolation Cytotoxicity Permeability Cell technology 

1 Prospects of Cell Encapsulation

Entrapping the biological cells in physical polymeric membrane is the heart of medical sciences for treating the ailments that have been a dream of the medical professionals for years. This technique has been introduced as a basic research tool in 1950s to trap the living cells in a nonliving membrane where the cell carries out its normal metabolic activities. In 1964, Chang introduced the term “artificial cell” for encapsulated cells [1]. Cell encapsulation is a technique where the transplanted cells are protected from the host’s immune system or tissue rejection by enveloping the cell in an artificial, partially permeable polymeric membrane which potentially allows the grafting of the cells without using immunosuppressant drugs [2]. This technique allows scientists selectively enter the medicine in the cells through semipermeable membrane. Cell encapsulation has the ability to cure the diseases including lysosomal storage disease (LSD) [3], neurological disorders, e.g., Parkinson’s disease [4] and dwarfism [5], and genetic disorder, e.g., hemophilia [6], diabetes [7], and cancer based on immune-isolation gene therapy [8, 9].

The encapsulated cell selectively absorb the essential nutrients that are required for its survival based on the differential permeability of the membrane, but the other molecules especially the large sized molecules cannot diffuse through the membrane [10]. Despite the nonliving nature of the membrane, it allows various cellular activities such as cell differentiation, metabolism, proliferation, and cellular morphogenesis while protecting the cell from the host immune system, this concept is called immunoisolation, depicted in Fig. 1. The main incentive in cell encapsulation is to incorporate the cells in the recipient’s body without using the immunosuppressant drugs [11]. These drugs are generally administered to suppress the immune response to the transplanted cells as these cells are crossing the recipient’s immunological barrier. The hindrance in using immunosuppressant drugs is their systemic administration, which would suppress homeostatic immunological physiology of the immune system leading to undesirable consequences including opportunistic infections, and it could adversely affect the incorporated encapsulated cells. Ethically it is not advisable to administer immunosuppressant drugs even to those patients for whom the organ transplant is needed. Therefore, it is the demand of the day to develop such methods which could be applied to the patients and treat disease without medicating the immunosuppressant [12].
Fig. 1

Immunoisolation of a cell encapsulated by hydrogel matrix [13]. (With kind permission of SciTech)

Cell encapsulation is a technique in which living cells are enveloped by a nonliving semipermeable membrane in such a way that the cells can grow and express their genes even after encapsulation by the nonliving membrane (Fig. 1) [13]. Since the cells are encapsulated by the biocompatible polymer based nonliving semipermeable membrane, the patient’s immune system is unable to detect the presence of encapsulated cells. Therefore, the key role of cell encapsulation is to transplant the cells into the host body without any disturbance in the normal homeostasis of the body.

Whenever the donor naked cells are directly enter into the body of the recipient, the immune system of the recipient is triggered, the host immune system activated and releases huge number of defensive cells, e.g., phagocytic cells and lymphocytes in response to the foreign cells. The cells of the host immune system start producing cytotoxic molecules and cytokines that destroy the transplanted cells. The situation intensified especially in those cases where a less number of donors are available for cell donation or the required donor cells are difficult to grow on artificial culture medium. Cell encapsulation provides an adequate solution to these problems faced by organ transplant experts and also helps relieving the patients suffering from serious complications [14]. Various means of cell encapsulation has been used for curing the human illnesses like direct transplantation of the cells. This process implants bovine adrenocortical cells (BACs) encapsulated by alginate and used as a replacement therapy for adrenocortical insufficiency. It is always desirable to implant the immune-isolated foreign cells directly into the patients. However, this technique is limited in its use in clinical applications because of the unavoidable risk of direct contact of the foreign cells to the host immune system. In addition, the degradation of the matrix could lead to the escape of matrix components and develop sensitization to the newly exposed cells [15]. This limitation is now overcome by using the hydrogel that not only protect the implanted cells but also take care of the nutrients entering the cells for their survival. In order to understand the concept of cell encapsulation, a schematic representation of cell-selective encapsulation is given in Fig. 1a. The experimental illustration of hydrogel formation around the cells shown in Fig. 1a is further elaborated in Fig. 1b where the cells are encapsulated in two steps: (i) antibody-antigen interaction for identification of cell and (ii) formation of hydrogel around the cell by a reaction catalyzed by horseradish peroxidase (HRP). Until now, hydrogels have been prepared from aqueous solution containing HRP of a variety of polymers including natural polysaccharides [16, 17], derivatives of polysaccharides [18, 19], proteins [20], and various synthetic polymers [21, 22].

Numerous polymeric materials are being explored for the encapsulation of therapeutic cells. Synthetic poly(ethylene glycol) (PEG)-based hydrogels are used in which gelation is done by photopolymerization using monomers, photoinitiators, and cross-linkers. Using microwave irradiation, PEG is chemically modified with methacrylic anhydride to get polyethylene glycol dimethacrylate (PEGDM), which was then exposed to 2-hydroxy-1-[4-(hydroxyethoxy)phenyl]-2-methyl-1-propanone photoinitiator in presence of ultraviolet light [23]. Cross linked alginate was used for the first time to encapsulate islets protecting them from immune system when transplanted in the body of the patient [24]. Initially liquid-core encapsulation method and syringe pump extrusion techniques were employed for encapsulating islets, but later on a more matured and modified approach was developed [25]. In this approach, the islets containing alginate droplets are poured into a solution containing calcium chloride for gelation. This technique is found to be applicable to a large number of cells, e.g., dopamine-secreting cells. The cells encapsulated in alginate are found to be stable in in vitro cultures for 4 months and survive in vivo when transplanted intraperitoneally in T1D human recipients [26, 27]. In spite of strenuous work by the scientists, they encountered some problems with the alginate beads that include nonuniformity, broad size distribution, and permeability of alginate beads. Wolters et al. [27] developed an air jet droplet generator system to address the existing problems of alginate beads. However, it was found that the beads produced by air jet droplet generator develop “tails” that causes immune responses. Moreover, the pressure applied during air jet operation causes damage to the cells. Sometimes air bubbles also get entrapped in alginate beads and hinder the diffusion of vital nutrients that is required for the survival of the cells [28].

Different materials follow different gelation mechanisms for the encapsulation of cells. In case of thermally sensitive gels, gelation mechanism may be controlled by temperature, e.g., gelation of gelatin and agarose. The gelation of photosensitive PEG is controlled by UV light [29]. However, the exposure of light and temperature could damage the cells during encapsulation. Therefore, the ion-based method is preferred in most of the cases for gelation purposes [30].

There are two types of polymers that used in cell encapsulation, namely (i) natural polymers such as alginate, agarose, collagen, and chitosan, and (ii) synthetic polymers such as PEG, poly(lactic-glycolic acid) (PLGA), poly(lactic acid) (PLA), and poly(glycolic acid (PGA). Among all of these polymers used in cell encapsulation, alginate is found to be the best material for applications in biological systems due to its easy crosslinking, biodegradability, and biocompatibility. Yet, alginate has a number of shortcomings that prevent them from using in in vivo experiments. One of the main drawbacks of alginate is cellular overgrowth that might be caused by immune response due to the release of foreign cellular material leaking through alginate membrane or to the exposure of the cells to the membrane due to the breakage of the cellular membrane. The other problems associated with the utilization of alginate capsules are that the alginate capsules are relatively weak, have increased pore size, and dissolve/destroy upon long term exposure to the host immune system. In order to increase the stability and maintain the porosity of the capsule, a polycation layer is usually added to the capsule, but most cases this polycation layer also cause inflammatory reaction to the host.

Another significance aspect of cell encapsulation is the drug delivery from the encapsulated cells. By adopting this system, it is possible to deliver/release the exact amount of drugs in a controlled manner that minimized side effects without repeating its use, which ultimately improve the quality of life [31]. Various types of polymers including both natural and synthetic polymers are being used for slow release of therapeutic agents, proteins, growth factors, and drugs. The drug delivery system was classically based on alginate-poly-l-lysine-alginate (APA) microcapsules. The main disadvantage associated with the APA microcapsules is their biodegradability. In order to address this challenging issue, the thickness of the membrane material must be optimized so that the capsule remains intact for the designated time period of use in the body of the patient. Ma et al. [32] have shown in his work that it is possible to optimize the thickness of the membrane which in turn varies the biodegradability of the APA capsules. For instance, an increase of 0.02% (w/v) to 0.08% (w/v) in the concentration of PLL solution causes an increase of 0.4 μm in membrane thickness of APA capsule. Conversely, an increase in the pH of PLL solution from 5.8 to 9.2 results in the decrease of the thickness of the membrane from 9.8 to 8.6 μm. Therefore, it is possible to attain required thickness and biocompatibility of APA membrane by changing the specific reaction conditions under which the membrane is synthesized [32].

Research scientists are moving a step forward developing microcapsules with low durability and remarkable biodegradability that includes poly(lactic acid) and its co-polymers with glycolic acid, lactides and glycolides and their copolymers, polyalkyl cyano acrylates, polyanhydrides, and corbopol. These polymers have been known to be biodegradable for many years. These newly designed 3D scaffolds without additional coating have the potential for cell delivery to the patient instead of just releasing the therapeutic agents [33]. Therefore, the release of drug is much easier and faster as compared to classical APA microcapsules. The rate of degradation of the microcapsule depends on the number of hydrolyzable functional groups present in the polymer composing the microcapsule wall. The increased number of hydrolyzable functional groups in the polymer backbone, the more will be the rate of degradation. Additionally, in certain circumstances the extremely high rate of degradation is required to release the cells, in these cases completely oxidized alginate-fibrin is used to accelerate the degradative process, e.g., injectable scaffolds in which calcium phosphate cement is used to release the cells quickly [34]. A comparison of degradative process and durability of the microcapsules is depicted in Fig. 2. However, drug delivery as well as cell delivery systems are becoming more and more attractive these days for curing serious illnesses, such as carcinomas especially brain tumors where the release of the cells or therapeutic agents are required directly into the specific cells [35]. In addition, the other encapsulated cells like encapsulated islets have been used as a therapy for curing diabetes saving the lives of a large number of patients across the world [27, 36]. Cell encapsulation could mimic the role played by gene therapy where the encapsulated cells containing gene of interest could be implanted directly into the body of the patient to release the therapeutic agents. Gene therapy is the technique to make up for the faulty genes to cure the illnesses.
Fig. 2

A schematic diagram showing a comparison between degradative process and durability of the microcapsules, i.e., both classical APA design and new scaffolds for drug delivery and cell delivery [35]. (With kind permission of Science Direct)

Gene therapy can be classified in two ways, i.e., in vivo and ex vivo gene therapy. In vivo gene therapy consists of all the methods of gene therapy whereby the gene of interest is directly transferred into the body of the patient [37]. In vivo technique is preferred when there is no option of isolating and growing the individual cells from the affected part of the body as in case of neural disorders [38]. On the other hand, in case of ex vivo gene therapy, the gene is transferred into the patient-derived cells grown in the artificial culture medium. These cells are then expanded by culturing in vitro and ultimately transferred into the patient where the genes express themselves and make the desired product of interest. This technique is used when the removal of the cells from the body is easier where the cells are genetically modified and engrafted back into the body of the patient [39].

The first use of agarose came on the scene when Nilsson et al. used agarose for encapsulating the islets of Langerhans [40, 41]. They showed that a large number of cells were encapsulated in agarose, and the microcapsules were found to be stable. Unfortunately, this work was not advanced further, but at a later stage it was found that this technique appeared to be potential for the preparation of artificial organs such as biohybrid artificial pancreas (BAP) that consists of living cells and biomaterials [42]. Considering the stability, nontoxicity, and mechanical strength, the agarose is found to be the suitable for encapsulating the mammalian cells. However, in vivo studies have shown that the effect of the agarose encapsulation decreases on the islets of Langerhans with the passage of time. In an experiment where the agarose encapsulated cells were implanted in the rat for normoglycemic activity, it was found that the activity of these cells has dropped to 50% in 15 days [43]. Although the agarose showed many promising and intriguing properties, there are many issues yet to be resolved concerning the encapsulated agarose islets of Langerhans before considering them for real time application [44]. As a whole, the success of the encapsulation technique depends on the properties of the encapsulated cells, their performance in the living organism, and the materials used in the matrices.

2 Materials for Cell Encapsulation

The materials used in cell encapsulation must be biocompatible and have reasonable mechanical strength to survive handling, implantation, and the mechanical, biochemical, and biological stresses imposed by the host. The capsules membrane must allow the exchange of oxygen, nutrients and metabolites, while obscuring the encapsulated cells from the host’s immune system. In the search for a better encapsulation design, many types of natural and synthetic polymers are being explored. Some of the most popular methods and materials, which have been applied to cell encapsulation, are shown in Table 1.
Table 1

Methods and materials used in cell encapsulation

Methods

Materials

Ref.

Bulk hydrogel

PEG-based hydrogel

Photopolymerizable styrenated gelatin

Silk fibroin

Photopolymerized elastin-like polypeptide

Chitosan chondrocytes

Photopolymerized hyaluronic acid

Collagen and laminin

d-mannitol crystals with photocrosslinkable MAC

Alginate di aldehyde (ADA), gelatin, borax

[45, 46]

[47]

[48]

[49]

[50]

[51]

[52]

[53]

[54, 55]

Hollow fibers

Poly(acrylonitrile-vinyl chloride) [PAN-PVC]

Chitosan, alginate in PAN-PVC

Polysulfone

Collagen in polysulfone

[56]

[57, 58]

[59, 60]

[61]

Beads

Alginate, gelatin

Alginate, poly-l-lysine (PLL)

Alginate, chitosan fibroblasts

Agarose

[62, 63]

[64, 65]

[66]

[67]

2.1 PEG-Based Hydrogels

Hydrogels are defined as 3-dimensionally cross-linked network of any profoundly water-soluble polymer. Additionally, hydrogels could be developed in different forms, e.g., slabs, micro- and nanoparticles, in the form a coating or a film, and are used for different clinical applications [68]. Polyethylene glycol (PEG) is a condensation polymer of ethylene oxide and water. PEG-based hydrogels are encapsulation systems that provide biocompatibility, long term stability, and permeability making PEG-based hydrogels ideal for protein release from the cells [45]. As the PEG has limited affinity for the cells and the biological molecules, its affinity can be increased by adding various cell binding peptides. The bioactivity of the PEG-based hydrogels can be increased by the addition of the cell binding peptides and cell modulators [69]. Though PEG is not considerably biodegradable, the different biodegradable components such as polyesters, fumarates, acetals, disulfides, and enzyme sensitive peptides are incorporated into PEG hydrogels enhancing the biodegradability of PEG-based hydrogels [70]. For example, dimethacrylated PEG (PEGDM) and PEG urethane dimethacrylates (PEGUDM) are synthesized by chemical reaction using hydroxyl groups containing PEG chains with methacrylic anhydride and 2-isocyanatoethyl methacrylate, respectively, in presence of tetraethyl amine (TEA). Previously, PEG hydrogels have been synthesized by interfacial photo-polymerization in which PEG was coated around a single cell by a photoinitiator that absorbed onto the surface of the cell and activated by argon ion laser [71]. It is worth to mention that the photo initiated polymerization of hydrogels was proven to be a compatible technique for encapsulation of many cell types [72].

These hydrogels are known to have tunable properties. It is possible to change the structural properties of the hydrogels (e.g., crosslinking density, degradability) without changing the properties of the polymer [73]. PEG is known to be a biocompatible material and helps improving the mechanical strength of other encapsulating materials. Many of the structural properties mainly transport properties of the hydrogels can be easily controlled by changing the concentration and controlling the molecular weight of PEG chains in the initial reaction mixture. In general, the control of the transport properties of the gel is highly critical to maintain the balance of the nutrients and other metabolites for the survival of the cells while protecting the cell from the immune system of the patient [74].

Oligo poly(ethylene glycol) fumarate (OPF) is a polyester that is synthesized by the condensation reaction between PEG and fumaryl chloride in the presence of TEA. OPF was found to be a support for the formation of tissue in bones, cartilages, osteochondral cells, tendons, cardiovascular, ocular, and neural tissue [75]. OPF hydrogels could be used to encapsulate numerous types of cells such as chondrocytes, embryonic and mesenchymal stem cells, and connective tissues like primary tendon fibroblasts. In addition, the OPF hydrogels can also be used to deliver growth factors, DNA plasmids, small molecule, nanoparticles, and chemotherapeutic agents [76]. The crosslinked OPF prepared by thermal initiator is used as an injectable, biodegradable capsule for transplanting and regenerating the bone tissues. In an experiment, the rat marrow stromal cells (MSCs) were linked with the hydrogel precursors and encapsulated by using OPF solution at a concentration of 14 million cells/mL. These cells were then cultured in vitro in a medium containing osteogenic supplements (dexamethasone). The growth of the cells was monitored continuously and the results obtained at day 7, 21, and 28. It was found that the mineralized matrix is not formed at day 7, but at day 21 the mineralized matrix become visible, and at day 28 the cells have generated mineralized matrix throughout the samples, which clearly indicate that thermally cross-linked OPF hydrogels are highly useful as an injectable cell carrier for bone regeneration [77].

Despite the above advantageous properties, the OPF hydrogels do have certain disadvantages. Firstly, low tensile strength of hydrogel makes their use limited to low load bearing applications, and the softness of the hydrogels can result in dissolution and removal of the hydrogel from the site of implantation. Secondly and most significantly, the problem is the release of drugs from hydrogels particularly in case of nonpolar and hydrophobic drugs, the loading of exact quantity of drugs while maintaining their homogeneity in hydrogels becomes difficult. Thirdly, rapid release of the drugs from the hydrogels due to high water content and large pore size [78].

2.2 Photopolymerizable Styrenated Gelatin

As we have seen in case of hydrogels, they suffer from the lack of considerable mechanical strength. Therefore, it becomes difficult to keep the cells fixed at a particular target site in the host [79]. Several biomaterials including natural and synthetic polymers have investigated for the improvement in the properties of these materials making them desirable biomaterials for cell encapsulation. An ideal cell encapsulating material should have four desirable properties: (a) it should have fix transplanted cells at the donor site, (b) should be biocompatible, (c) must have moderate biodegradability in order to avoid complications during post-transplantation period allowing gradual replacement of the cells from the body by the newly synthesized tissues, and (d) the biomechanical properties of this material should be similar to those of normal tissues of the host [47, 80]. Unfortunately, such material has not been developed until now, but continuous efforts are being made in this regard to synthesize a new material having these above mentioned properties.

The photo-polymerized hydrogels could be easily crosslinked by means of light-induced polymerization of a monomer or a macromer and achieve desired properties [81]. One of the salient features of photopolymerizable hydrogel is the formation of crosslinked hydrogel around the transplanted cells which could be carried out in situ from injectable solution and makes these hydrogels attractive for use in medicine. A huge number of research studies have been conducted that support the use of photopolymerizable hydrogel as scaffolds for culturing chondrocytes [47, 82, 83]. Accordingly, photopolymerizable styrenated gelatin have been developed that can undergo gelation simply by the photopolymerization induced by the visible light. A schematic representation for the synthesis of photopolymerizable styrenated gelatin is shown in Fig. 3.
Fig. 3

Schematic diagram showing photopolymerization of styrenated gelatin chains synthesized by using camphorquinone as photoinitiator. When visible light shines on CQs, free radicals are generated and polymerization of the styrene group is initiated [47]. (With kind permission of Mary Ann Liebert Inc.)

The styrenated gelatin is synthesized by the reaction between the hydroxyl group of 4-vinylbenzoic acid and amino group of lysine residue of gelatin. These styrenated gelatin (95 kDa) chains are then used for gelation in two steps. In first step, the camphorquinone (CQ), a photoinitiator, undergoes hydrogens abstraction as it is exposed to irradiation by visible light and produces free radicals, and in the second step of initiation, it combines with styrene group and polymerization takes place [84].

Photopolymerizable styrenated gelatin has many advantages. It can tolerate huge load, show good cell viability, and appear as a viscoelastic biomaterial with elastic modulus of 13.41 kPa. These gelatin hydrogel has a great potential to be used as a sustainable biomaterial for biomedical applications. The significant advantage of the styrenated gelatin is its ability to undergo in situ gelation from injectable solution making it a desirable biomaterial for several clinical applications such as cartilage tissue engineering [85], adipose tissue culturing [86], and anticytokine antibody therapy in surgically restricted area [87]. In spite of the abovementioned advantages, there is still roam to modify the photopolymerizable styrenated gelatin to enhance the cell viability and deserve optimization of the construction of the matrix while transplanted into the host [36].

2.3 Photopolymerized Elastin-Like Polypeptides (ELPs)

Elastin is a protein that is found in the extracellular matrix especially in connective tissues, i.e., blood vessels, skin, and ligaments [88]. Elastin is synthesized from its soluble precursor protein called tropoelastin, a protein consisting of hydrophobic and hydrophilic crosslinking domains. When tropoelastin is secreted into the extracellular space, it starts crosslinking through lysine residues resulting in the formation of insoluble elastin protein. Keeping the role of elastin in extracellular matrix (ECM), it is becoming an interesting area in applications of drug delivery and tissue engineering [89]. Elastin-like polypeptides (ELPs) consists of pentapeptide repeats derived from the hydrophobic domains of tropoelastin. The most commonly used hydrophobic repeat in ELPs is (VPGXG)n, where X represents “guest residue,” i.e., any amino acid other than proline, and n represents the number of pentapeptide repeats in ELP. ELPs consisting of (VPGXG)n sequence show a very unique property called inverse temperature phase transition. The ELPs are soluble below or above their transition temperature. Upon demixing of ELPs, they become insoluble in aqueous medium and aggregation of ELPs occurs into its structures called “coacervate” phase [52, 90]. The reversibility of transition temperature could be easily tuned by changing the composition of the guest residues and the length of the ELP chains. In addition to inverse temperature phase transition, ELPs also show response to stimuli and thereby belong to a stimulus responsive class of polymers called “smart polymers” [91]. These polymers show response to the changes in the environment of their solution. ELPs are not only responsive to the changes in temperature but also show response to change in pH, redox trigger, and light. Because of these tunable properties, ELPs are useful biological polymers that show response not only to the temperature but also other environmental factors such as pH and ionic strength. In addition, as ELPs are composed of amino acid sequences, it is possible to control the size of the ELP by controlling the size of the encoding gene [51, 92, 93].

A very precise control over size, stimulus response, and sequence of amino acids in ELPs make them an interesting platform to design drug carries for both local and systemic applications [94]. Tissue engineering of the ELPs allow the incorporation of the amino acids of interest into ELPs which make them targeting protein like a ligand to the receptor on the cell surface and promoting the uptake of the pharmaceutical compound into the cell. Because of the control on number and sequence of amino acids, it is possible to design the drug binding sites in ELPs for conjugation of the drugs or any locating probe [51, 95]. Because their biocompatibility, genetically encoded synthesis, and responsiveness to the stimulus are highly tunable properties, ELPs are becoming extremely attractive as macromolecular carriers of drugs for curing cancer [51]. By precisely exploiting the aforementioned properties, ELPs have been synthesized incorporating different proportion of valine, glycine, and alanine with pentapeptide repeats and transition temperature ranging in between 37 °C and 42 °C which could be obtained by focused ultrasound or microwave on the specific parts of the body. The ELP was then conjugated with rhodamine, injected intravenously, and monitored its location and accumulation providing local and slight hyperthermia to the tumor [51, 96].

ELPs are attractive biomaterials for encapsulating the cells. They are derived from extra cellular matrix (ECM), and they could provide ECM-like environment to the cells and tissues by providing the factors similar to their natural environment [93, 97]. Availability of highly tunable properties such as control of polymer chain length, composition, sequence of amino acids, ease of making different polymer architecture and type, number and location of crosslinking sites makes ELPs highly compatible for cell encapsulation. ELPs form triblock polymer by physical crosslinking in which the central block is of hydrophobic component and two hydrophilic blocks on two sides of the central block. The hydrophobic portion forms the gel above the transition temperature by hydrophobic interactions, and the strength of these gels depends upon the length of the hydrophobic chain and solvent. Greater hydrophobic content allows higher elastic modulus. However, in some cases much greater strength is required, e.g., load bearing applications in case of cartilage repair, which could be sorted out by chemical crosslinking. Chemical crosslinking of ELPs could be accomplished easily in aqueous medium using hydroxymethylphosphines (HMPs). ELP hydrogel was formed through crosslinking between ELP chains and lysine residues of [tris(hyroxymethyl) phosphine] propionic acid (THPP). The hydroxyl group of THPP reacts with amino group of lysine residue in ELP chain resulting into the formation of hydrogel as a function of pH. The mechanical strength of the hydrogel could be controlled by controlling the number of lysine residues in ELPs chain [98].

2.4 Photopolymerized HA, Collagen, and Laminin

Hyaluronic acid (HA) is a biopolymer discovered by Karl Meyer and capitalized by group of researchers in the vitreous of bovine eye in 1934 [51, 99]. HA is a naturally occurring biopolymer found more abundantly in the connective tissues especially in the joint synovial fluid, vitreous fluid in eye, and umbilical cord [100]. It is naturally synthesized in the cells by special membrane proteins called hyaluronan synthases and is degraded by a group of enzymes called hyaluronidases. HA is composed of a disaccharide repeating unit consisting of d-glucuronic acid and d-N-acetylglucosamine linked together through β-(1-4) linkages [101]. Both sugar units resemble glucose, which in β position allows its bulky groups and anomeric carbon to be at the less sterically hindered equatorial position and smaller atoms like hydrogen to be at less favorable axial position giving HA a conformationally stable structure.

HA is a useful biological molecule. It plays an important role in many biological processes such as cell proliferation, wound healing, morphogenesis, inflammation, and migration [102]. In addition, hyaluronidase is an enzyme that could easily degrade HA to mark HA a biodegradable polymer. Based on these properties, HA could be considered as an excellent fabricating materials that could serve as an extracellular matrix for encapsulating the cells. In recent past, it was found that the photo-crosslinked HA hydrogels are excellent scaffolds for heart valve and cartilage tissues engineering [103].

In a recent work reported by Bae et al., HA hydrogel beads have been utilized for encapsulating bovine articular chondrocyte using a novel technique [104]. HA hydrogel beads were prepared by photopolymerization as shown in Fig. 4. In this process, presynthesized methacrylate-HA was added to the solution containing calcium chloride, triethylamine, alginate, and N-vinylpyrrolidone. 4-Benzoylbenzyltrimethylammonium chloride was used as a photoinitiator. In order to carry out the photopolymerization of methacrylated HA and N-vinylpyrrolidone, long wavelength UV light was used for 60 min in nitrogen atmosphere. The resulting beads were incubated in EDTA solution for a day to extract alginate and then washed extensively with deionized water.
Fig. 4

Schematic representation for the preparation of HA hydrogel beads

The HA hydrogel beads (Fig. 4) were then used for encapsulation of bovine articular chondrocytes by microinjection technique in which the cells were directly injected into the beads by using a very fine needle with a tip diameter of 248 μm. An injection needle defect was created in the beads which was sealed by adding poly-l-lysine solution to the encapsulated beads to avoid the slipping out of the cells from the beads. The bovine articular chondrocyte encapsulated in this way was able to proliferate and survive in the beads (Fig. 5).
Fig. 5

Schematic representation of cell encapsulation by direct injection using microinjection technique

Collagen is another important natural protein of ECM and has been extensively used for encapsulation of cells as extracellular matrix. Collagen can be used in many forms including soluble collagen hydrogel, collagen sponges, fibrillary collagen cross-linked with glutaraldehyde. Collagen hollow fibers and microcapsules have been successfully used to encapsulate the artificial liver to encapsulate hepatocytes [105]. Though the naturally derived collagen has been used in many applications including bone repair, tissue engineering, and drug delivery systems, but it has some disadvantages like difficulty in modifying the sequences of the naturally derived collagen and possibility of immunogenic response [106]. In order to overcome these issues, collagen peptides have been chemically synthesized as an alternative biomaterial with improved properties. Since it is chemically synthesized, it is possible to control the amino acid sequence according to the requirement for providing the controllable materials for matrix engineering [107]. Recently, the covalently crosslinked collagen hydrogel has been synthesized by an 8-arm poly(ethylene glycol) star polymer with collagen-based peptide, which results a viscoelastic thermoresponsive hydrogel in nature. These crosslinked hydrogel is appeared to be with desirable properties such as stiffness and highly cross-linked network of pores where cells could easily reside while turning these hydrogels into promising scaffolds for cell encapsulation [60, 108].

Laminin is also an integral component of the neural tissues ECM that composed of three chains, and found in excess on the inner face of endoneurium where it is found in close contact with the Schwan cells and neurons [109]. The role of laminin in cell encapsulation is found to be of considerable impact with collagen and HA hydrogels. Recently, a study carried out by Suri et al. has shown that 3D interpenetrating network (IPN) hydrogels prepared from collagen and HA were used for encapsulating the Schwan cells [110]. Two types of hydrogels were prepared: one containing both collagen and HA in photocrosslinked form called IPNs and other containing only collagen in photocrosslinked form, and HA fibers are just entangled in the collagen network called semi-INPs. A schematic representation of synthesis of IPNs and SIPNs is given in Fig. 6. The synthesis of IPNs and SIPNs begins by suspending glycidyl methacrylate modified hyaluronic acid (GMHA) and collagen solution in a silicon mold where collagen is allowed to undergo fibrillogenesis at 37 °C labeled as (1) in Fig. 6, and this allows the formation of SIPN which upon exposure to UV light results in photocrosslinking of GMHA leading to the formation of full IPN (2). The step (3) illustrates the effect of photo masking which allows photo patterning of the hydrogel leading to IPN and SIPN in the same bulk hydrogel. Both IPN and SIPN were applied as novel devices for neural regeneration therapies. It was found that the Schwan cells surrounded by 3D hydrogels containing laminin were not only able to survive for 2 weeks but also proliferated and carried out their secretary role to enhance the growth of the neurons. Overall, all these three components, namely, collagen, HA, and laminin, could mimic the role of extracellular matrix and could be used to encapsulate the Schwan cells that may serve as a nerve regeneration therapy. It may further enable us to understand the basis of Schwan cell interactions with neurons and different ECM components [111].
Fig. 6

Schematic representation of synthesis of IPN and SIPN [110]. (With kind permission of Science Direct)

2.5 Silk Fibroin

The silk is considered as one of the best biomaterials due to its highly significant properties such as biocompatibility, easy to be chemically modified, gradual rate of in vivo degradation, and its potential to be changed into different forms of biomaterial either in aqueous solution or organic solvent [112]. The silk is collected from silkworms especially Bombyx mori. The raw silk is obtained from cocoons, which have two components, namely, sericin and the core protein called fibroin. The sericin component is removed from the raw material and core fibroin is obtained, which is used for different biological applications. The sericin is composed of soluble glycoproteins. They are sticky in nature. The sericin completely surrounds the core fibroin protein of cocoon filament. Once the sticky sericin is removed, the remaining component is called silk fibroin [41, 113]. This silk fibroin showed great potentials to be used in the synthesis of different biomaterials shown in Fig. 7.
Fig. 7

Different formats of biomaterials derived from silk fibroin [112]. (With kind permission of Springer Nature)

The silk fibroin consists of two chains: high molecular weight (MW) (Mw ∼ 390 kDa) and low MW (Mw ~ 26 kDa). They linked together through disulfide bridges. The silk fibroin is a block copolymer and consists of hydrophobic β-sheet forming blocks which are linked together through hydrophilic cross linkers [41, 114]. The crystalline portions of silk fibroin consist of glycine-X repeats, where X represents serine, alanine, threonine, or valine amino acids, and in between these repeats they linked with other domains that consist of glycine, alanine, serine, and tyrosine residues. This arrangement of amino acids gives rise to a resiliently strong protein. The strength and toughness of this protein is further enhanced by the presence of β-sheets [115]. The toughness of silk fibroin is found to be much better than the best synthetic biomaterials, Kevlar. The strength of silk fibroin is even better than commonly used degradable polymeric biomaterials especially collagen and poly(l-lactic acid) (PLA). A comparison of ultimate tensile strength of collagen (0.9–7.4 MPa) and PLA (28–50 MPa) with that of silk fibroin (740 MPa) shows that silk fibroin is far better in strength than the other two materials used for various applications [41, 116].

The silk fibroin is used for various applications at cell level. It has been found in an in vitro experiment that an aqueous solution of silk fibroin self-assembles and forms β-sheets leading to the formation of hydrogel. This sol-gel transition of silk fibroin is influenced by variation in temperature, pH, and ionic strength [117]. It is also found that the strength of the silk fibroin increases by increasing the concentration of silk fibroin in the aqueous solution. For cell level applications, it is always desirable that gelation may occur at a faster rate, i.e., in hours. However, in case of silk fibroin, the gelation takes place at longer time unless some modifications are executed in the natural silk fibroin. For example, it is seen that in 0.6–15% (w/v) solution of silk fibroin, sol-gel transition occurs in days or even weeks at 37 °C. Addition of salts has shown no influence on reducing the time of gelation, but lowering of pH decreased the time of gelation to few hours. These modifications have some effect on decreasing the time required for gelation, but they could compromise the normal physiological activities of the cell including their viability [45, 118].

In order to resolve this issue, a new method called ultra-sonication has been developed to accelerate the sol-gel transition of silk fibroin in a controlled manner. The gelation could be controlled by changing the power output and sonication time. This technique is based on the gelation by the formation of β-sheets due to change in the hydration of the hydrophobic component. In order to encapsulate cells or drugs in the silk fibroin hydrogels, the sonication conditions are optimized first without the cells, and the cells are then added before gelation and maintain a homogeneous and favorable medium that is required for cell encapsulation. For example, the cells have been encapsulated by autoclaving 4% (w/v) solution of silk fibroin. In another case, the human mesenchyme cells have been encapsulated by adding 50 μL of cell suspension to the sonicated silk at 37 °C. It is found that low concentration of K+ ions and lower pH accelerates the rate of gelation, but the presence of Ca++ ions and higher concentrations of K+ ions reduces the rate of gelation. Nevertheless, the cells encapsulated by sonication of silk fibroin retained their physiological activity and proliferation over weeks [45, 119].

2.6 Chitosan

Chitosan is a derivative of chitin. When the deacetylation of chitin reaches to 50%, it becomes soluble in aqueous acidic medium. The reason for solubilization is the protonation of –NH2 group at C2 of the d-glucosamine repeating unit in chitosan and appeared as a polyelectrolyte in aqueous acidic medium [120]. The degree of deacetylation can be determined by NMR spectroscopy, and the functionality of chitosan could be changed by changing the degree of deacetylation [121]. Chitosan is extensively used in controlled and slow release drugs due to its highly desirable properties that include nontoxicity, ease of biodegradability, and gel forming ability at low pH. Additionally, chitosan usually does not pose any stomach irritation as it has antacid and antiulcer activities; the formulations made from chitosan float and gradually increase in volume in acidic medium [122].

Chitosan has been used as a polymer matrix for cell encapsulation for curing Parkinson’s disease, hepatocytes, human bone marrow cells, and cardiomyocytes [123]. It is commonly used for cells that grow favorably in cationic environment or for biodegradable applications. However, chitosan has not been explored as much as agarose or alginate due to the fact that the chitosan is not easily dissolved in aqueous medium at pH greater than 6 except low molecular weight samples [124]. Nonetheless, chitosan is soluble at very low pH, that is, nonphysiological, and could turn out to be cytotoxic [125]. However, alginate-chitosan matrices have been successfully used for encapsulation and implantation of pancreatic islets in streptozotacin (STZ) induced diabetic mice [126]. But it is still not clear how the toxicity was avoided in this application. Actually, linking alginate with chitosan might improve both its biological and mechanical properties. Alginate and chitosan solution form hydrogel successfully at 37 °C when glycerol salt is added into the solution and has been applied for microencapsulation of insulinoma cells [127]. In certain instances, researchers have used N-acetylated chitosan instead of chitosan owing to its high solubility in the aqueous medium, but solubility decreases with increasing MW of the polymer [128]. The addition of alginate with N-acetylated chitosan resulted into a stable matrix for encapsulating the HPG2 cells that survived for 1 week [129]. Due to low compatibility of the chitosan and the encapsulating cells, it might not be accepted as a strong candidate for cell encapsulation, but it could be considered as a potential coating material [130].

2.7 d-Mannitol Crystals with Photocrosslinkable MAC

Another aspect of considerable importance prioritized while encapsulating the cells especially the neural cells is the porosity of the 3D scaffolds. The size and dimensions of the pore are factors of prime importance because they regulate the transport of oxygen, nutrients, and other cell surviving signals. The materials that produce pores on heating or dissolution are termed as porogen, and these materials are added into the scaffolds to develop porosity [131]. The dissolution could be carried out both in vivo and in vitro to check whether the porogen is biocompatible, does not alter the cell microenvironment, e.g., pH, osmolality, and most importantly it is eliminated from the scaffolds after developing pores. d-Mannitol is an interesting biomaterial in this class as it is biocompatible, easy to dissolve, and nontoxic to the cells, it has been used in bone cement and bone scaffolds [64, 132].

Recently, the methacrylamide chitosan (MAC) hydrogel derived from chitosan is used for Neural Stem/Progenitor cells (NSPCs) as a highly tunable growth matrix that offers differentiation, proliferation, and cell migration. d-Mannitol is used in this study to control the porosity of the hydrogel according to the requirements of the cells [133]. In this technique, MAC hydrogel is chemically synthesized by photopolymerization MAC in the presence of UV light by using IRG-184 (1-Hydroxy-cyclohexyl-phenyl-ketone) photoinitiator for encapsulating NSPCs and then d-mannitol crystals were added at different concentrations to the polymerizing solution. 20 wt% d-mannitol was found to be the best for optimized pores in the 3D scaffolds for NSPCs. Overall, it is found that d-mannitol along with photopolymerized MAC hydrogel is an excellent biomaterial for controlling the porosity of the scaffolds used in neural cell encapsulation [134].

2.8 Mixture-Induced Two Component Hydrogels (MITCHs)

A new type of hydrogel called physical hydrogel has been developed by exploiting the physical aspects of the polymer whose viscoelastic properties are tunable by molecular level engineering [135]. Such physical hydrogels usually formed without any external environmental triggers are injectable and highly compatible for use in cell encapsulation and drug delivery systems [136].

The previously known physical hydrogels were synthesized by changing the external environmental stimuli such as pH, temperature, and concentration of cations [137]. The main disadvantage of this physical hydrogel was the mixing of the cells with the precursors in the solution phase in which the cells were momentarily exposed to nonphysiological conditions which was not only detrimental for the cells but also for the proteins inside them. In order to get rid of these issues, a new strategy was developed to exploit the specific molecular interaction between two naturally occurring peptide domains to form two-component system, which has the ability to hetero-assemble into a hydrogel upon mixing under the constant physiological conditions, called the Mixture-Induced Two Component Hydrogels (MITCHs) [138]. By altering the sequence of amino acids in the primary structure, the binding affinity and the association domains can be tuned, and the physical hydrogels with highly controllable properties can be synthesized [139]. Figure 8 shows that there are two WW domains (CC43 and Nedd4.3) which have the ability to bind the proline-rich peptide (PPxY) mixed with hydrophilic spacers, and three polymer families are formed, i.e., C(x + 2), N(y + 2), and P(z + 2). A MITCHs hydrogel is formed when component 1 (either C(x + 2) or N(y + 2)) and component 2 P(z + 2) are mixed together at constant physiological conditions [140]. Three dimensional MITCHs hydrogels have been successfully used for encapsulation of rat mesenchymal cells, neural stem cells, and human endothelial cells [141]. Additionally, MITCHs also allowed the rat neural stem cells to differentiate into glial and neuronal phenotypes [142].
Fig. 8

There are two WW domains (CC43 and Nedd4.3) which have the ability to bind the proline-rich peptide (PPxY) and after mixing with hydrophilic spacers three polymer families are formed, i.e., C(x + 2), N(y + 2) and P(z + 2) [140]. (With kind permission of PMC)

2.9 ADA, Gelatin, and Borax

It is always desirable that the injectable polymer scaffolds used for cell encapsulation and drug delivery should be biocompatible and biodegradable [143]. The simplest approach is to directly inject the encapsulated cell or drug into the patient. The injectable polymer scaffolds provide ease of application, confined drug delivery for site specific action, better relief, and comfort for the patients [144]. Various injectable scaffolds have been prepared that includes pH sensitive, water soluble, temperature sensitive, and photopolymerizable hydrogels [145]. Hydrogels prepared from naturally derived polymers mimic the role of natural extracellular environment of the cells, and they cause the cell growth, differentiation and cell migration as well as stabilization of the encapsulated cells [146].

Many of these hydrogels are biodegradable and biocompatible, but they have certain disadvantages. These hydrogels are synthesized in situ using different crosslinking agents like enzymatic cross linker, chemical cross linker (such as glutaraldehyde, carbodiimide, adipic dihydrazide), and metal ions [147]. The photopolymerization requires a photosensitizer and longer duration of irradiation which is considered to be a limiting factor. The crosslinking by means of metal ions is reversible in the body and exerts cytotoxic effects [148]. The crosslinking agents that are incorporated into the hydrogel such as glutaraldehyde, polyepoxide, and isocyanate are highly toxic as they are gradually released while hydrogel degrade [149]. There are some crosslinking agents such as acyl azide and carbodiimide, which activates the crosslinking reaction without incorporation into the hydrogel, and are less toxic, but they slow down the crosslinking process [150].

Alginate is an anionic linear polysaccharide derived from brown sea weeds that constituents of 1,4-linked β-d-mannuronic (M) acid and 1,4-linked α-l-guluronic (G) acid residues [55, 151]. A very important feature of alginate is its gelation in the presence of divalent cations, e.g., calcium ions. Alginate has found wide use in cell encapsulation. However, alginate hydrogels degrade in an uncontrolled manner releasing calcium ions, and different sized alginate fragments exposed to the host [152]. High calcium level has also shown to inhibit the growth of cells in culture medium. It has also been found that high molecular weight alginate is not biodegradable, but its di-aldehyde derivative is easily biodegradable [153].

Alginate di-aldehyde (ADA) and gelatin has been used to synthesize injectable, in situ forming, and biodegradable scaffolds in presence of borax. In this hydrogel, both gelatin and borax are found to be biocompatible and safe to use in biomedical applications. Gelatin has a long history of use as a wound dressing material, and borax has long been used as a lethal dose in human beings [154]. The gelation reaction between alginate di-aldehyde and gelation in the presence of borax would take place without any extraneous force or factor like change in pH or temperature. It is further convenient to carry out under constant physiological conditions, and hence, the cells would not be exposed to non-physiological conditions at any stage of cell encapsulation. Furthermore, there is no threat of releasing divalent ions like calcium ions, and therefore, it would not affect the viability of the cells.

The ADA-gelatin system is not only suitable for cell encapsulation but also for controlled release of drugs for curing diseases [155]. Primaquine has been used to conjugate with ADA- gelatin hydrogel and its release was monitored. It was found that the extent of drug release depends upon the degree of gelation. At lower gelation, the release would be faster, but it becomes slower at higher concentration of the gel. Overall, there is much potential in the hydrogels to be used in various biomedical applications including cell encapsulation and controlled release of the drugs according to one’s desire for effectively treating the diseases. Now researchers are trying to make highly efficient systems for cell encapsulation that are biodegradable, nontoxic, injectable, and undergoing gelation under the normal physiological environment of the body to facilitate and provide great relief to the patients.

2.10 PAN-PVC

Poly(acrylonitrile-vinylchloride) (PAN-PVC) is a thermoplastic polymeric material, which is extensively used to fabricate the cell encapsulating hollow fiber membrane (HFM) [56, 156]. The HFM is one of the many important biomaterials originally fabricated for ultrafiltration applications [157]. PAN-PVC-based HFMs are well known for their transport properties allowing sufficient exchange of soluble factors for the viability of cells for extended period of time [158]. Previously an enormous amount of works focused on transport of the biological molecules based on molecular weight cutoff (MWCO) value and the molecular size at which 90% of the molecular species are rejected under hydraulic pressure [159], but it is established that HFMs transport properties influence biomass, proliferation, and controlled release of therapeutic agents from encapsulated cells [160]. HFM prepared from PAN-PVC using various solvents including dimethylformamide (DMF), dimethylacetamide (DMAC), dimethylsulfoxide (DMSO), and different additives and was formed to be different wall architecture. The HFMs formed from DMF, DMAC, and DMSO have finger like macrovoid structure as shown in Fig. 9ac, whereas the HFMs produced from additives are more symmetric with spongy architecture shown in Fig. 9d, e [161]. When the permeability of these HFMs was compared, it became evident that the HFMs prepared from DMF were more restrictive to the molecular diffusivity, while the HFMs produced from PEC additive were more permeable to the molecules for the growth of the cells. It was found that the permeability of PAN-PVC membrane increases irreversibly by washing with ethanol as compared to water [58].
Fig. 9

(ac) PAN-PVC HFM represented by showing SEM micrograph formed by using different solvents and additives (a) DMF, (b) DMAC, (c) DMSO, and (d, e) showing the wall architecture using DMF and different additives: (d) Dextrose, (e) PEO [161]. (With kind permission of Elsevier)

The PAN-PVC HFM system was also examined for its ability to release the drugs in a controlled manner. This was achieved by encapsulating the PC 12 cells that have the neurosecretory function and have the ability to release dopamine, a neurotransmitter [162]. The cells were cultured for 4 weeks and the effect of different membrane diffusivities was tested. It became very clear from the experiments that the cells were not only viable but also consistently released dopamine under different conditions for 4 weeks. However, it is still unknown that whether this treatment would be stable over a long period of time or not. In addition, the dopamine release also continuously increases as the permeability of the cells increases. The cells encapsulated by PAN-PVC membrane are highly permeable to the solutes from external environment while protecting themselves from immune system [57]. Therefore, by controlling the diffusivity of the PAN-PVC HFM, it is quite possible to make the release of the dopamine much controlled and regulated which indicating the potential of HFM for controlled release drug delivery.

2.11 Polysulfone

Polysulfone is prepared by condensation polymerization of bisphenol A and dichlorodiphenyl sulfone [59, 163]. The material is an amorphous high performance thermoplastic with a glass transition temperature of approximately 190 °C [164]. Polysulfones provide the advantage of preparing semipermeable membranes with developed inner and outer walls [60]. Usually the membranes including polysulfone membranes are prepared by phase inversion process in which a homogeneous polymeric solution is cast onto the support as a thin film and then immersed into a bath solution containing a nonsolvent [165]. The polymeric film solidifies to form a membrane with symmetric or asymmetric structure as a result of exchange between the solvent inside and the nonsolvent outside the cast film. The asymmetric membrane has an outer dense layer and a porous sublayer. These membranes have been widely used in liquid and gas phase separation as the thin top layer acts as a selective barrier and the sublayer which is composed of macrovoids and micropores provides good mechanical support.

The enormous amounts of research have been conducted in the past to understand the principle involved in the formation of membrane by phase inversion process. Chakrabarty et al. have proposed two types of demixing processes that take place during phase inversion, i.e., instantaneous demixing and delayed demixing [166]. According to this model, the membranes formed by instantaneous demixing generally show a highly porous structure with macrovoids and a finely porous thin skin layer. The membranes formed by delayed demixing show a macrovoids free and thick, dense layer [167, 168]. In order to get the membrane with desired porosity and structure, various additives are added during the fabrication of the membranes. Actually, the additives change the properties of the membrane by changing the phase kinetics or changing the solvent capacity and the thermodynamic properties of the membrane [169]. Various additive systems have been identified to play role in membranes like glycerol in a system of polysulfone, dimethylacetamide (DMAC)/water, maleic acid in a system of cellulose, dioxane/water/polyvinylpyrrolidone in a system of polysulfone [170]. Some additives have the tendency to form macrovoids, while the others suppress the tendency to form the macrovoids, thereby increasing the interconnectivities to increase the porosity of the top and sublayer.

A study conducted by Boom et al. showed that the addition of PVP additive in the formation of poly(ether sulfone) (PES) using NMP as a solvent decreases the possibility of formation of macrovoids by reducing the possibility of delayed demixing [171]. In another study, Yeo et al. reported that the addition of PVP to the polysulfone membrane using DMF as a solvent increases the possibility of macrovoid formation instead of suppression [172]. Jung et al. reported that the macrovoid formation depends upon the MW of PVP [173]. The addition of different MW of poly(ethylene glycol) such as PEG 400, PEG 6000, and PEG 20,000 Da showed a significant effect on the performance and morphology of polysulfone membrane prepared from the solvent NMP and DMAc separately [174]. Recently, polysulfone membranes have been constructed with the developed inner and outer surfaces. Such polysulfone membranes are obtained by means of special spinneret with appropriate conditions of hollow fiber spinning [175]. The membranes developed in this manner have found applications in biomedical sciences especially in the cultivation of the cells, e.g., hepatocytes [176].

2.12 Agarose

It is a polysaccharide which is used in cell encapsulation for many decades. It could form insoluble membrane when polyvalent cations are added to its aqueous solution [177]. Agarose is a gelling material which could be changed into thermally reversible gel by heating and cooling the aqueous solution of agarose [67]. Agarose is different from other polysaccharides. It could be changed into gel (~40 °C) below its gel-melting temperature (~90 °C). Agarose is stable at room temperature, and its gelling temperature lower than 37 °C is considered to be favorable [44, 178]. The bacterial strains could digest agarose, but mammals cannot. In an experiment, it was found that the cells encapsulated by agarose microcapsule can survive for 200 days when implanted in the peritoneal cavity of the rat. Therefore, the cells surrounded by agarose microcapsules are quite stable to be implanted in the human body [179].

3 Encapsulation Techniques

Cell encapsulation techniques have been broadly classified as macroencapsulation and microencapsulation. In macroencapsulation, a large group of cells are entrapped inside the hollow devices usually tube shaped or disc like, whereas in microencapsulation smaller groups of cells are entrapped in a capsule. The preferability of the technique depends upon the properties of the encapsulated cells. The problems that arose in macroencapsulation are resolved by adopting microencapsulation, which provide better control of physical and chemical properties of the encapsulated cells. The capsules with the size ranging from 0.3 to 1.5 mm are usually called microcapsules [180, 181]. Their small size as compared to macrocapsules increases surface area to volume ratio for transport of solutes across the membrane. Microcapsules are more durable and mechanically stronger than macrocapsules. Microencapsulation is carried out very carefully as the technique works on microscale and cells encapsulated in hydrogels undergo gelation and ultimately change into solidified matrix.

3.1 Macroencapsulation

As discussed earlier, macroencapsulation is a process whereby cells are encapsulated in the form of large groups as compared to microencapsulation. The cells are entrapped in the hollow devices having semipermeable membranes. The homeostatic working of macroencapsulating devices depends upon the homeostatic environment of the host. In general, the transport of the materials in or out of the cells depends upon the concentration gradient. The membranes of the macrocapsules are made up of thermoplastic materials having unique properties such as structural, functional, and mechanical properties [182]. The macroencapsulation technique was basically developed in the laboratory for understanding the concepts associated with transplantation of tissues and the immunology associated with transplanted tissues. However, macroencapsulation became a tool in combating ailments associated with endocrine system, nervous system, and metabolic disorders [183]. Now many companies worldwide are putting their full expertise to meet the challenges in macroencapsulation technology by generating the following modern devices/properties for producing products at a substantial scale.

3.1.1 Diffusion Chambers

This device is developed to understand the transport properties such as endocytosis, exocytosis, electrical conductivity, and resistance. The chamber uses the cultured cells for understanding the cell permeability of the cell membrane. The electrical conductivity of the membrane is measured by using the electrodes in diffusion chambers. The components of the diffusion chambers consist of cell culture device, diffusion chamber, and microreference electrodes. Cell culture device is 12 mm in diameter and 0.4 μm pore-sized polycarbonates and polyester membranes. The membranes are placed in the cell culture device using detachable rings. The cell culture device allows the placement of cells on the selectively permeable membrane. Thus, the cells are formed into a monolayer placed in contact with selectively permeable membrane. The next step is to introduce the cells in the diffusion chamber that may be vertical or horizontal. The monolayer obtained above is placed in diffusion chamber, and the transport properties of cells are studied [184]. By mid-1970s the diffusion chamber became one of the most advanced and unique device for studying cell differentiation and cell growth in the host. Many of the cell properties such as morphogenesis, cell division, differentiation, and release of chemical compounds from the implanted cells remain unchanged while using the diffusion chamber. The earlier membranes used in diffusion chamber were common macroporous membranes with pore-size of 0.10, 0.22, or 0.45 μm. Later nylon reinforced mixed cellulose esters were used in the membranes to maintain very fine and uniform pore size [185]. Such membranes are also used for various sterilization techniques to remove the germs from the desired products.

3.1.2 Ultrafiltration Membranes

The newer membranes were developed for the better performance of the encapsulated cells in the human body. For the first time, the ultrafiltration membranes were prepared using phase inversion process in 1960s. These membranes are commercially available in the form of flat sheets or hollow fibers. The most commonly used material for synthesizing ultrafiltration membrane is poly(vinyl chloride-acrylonitrile) (PAN-PVC) copolymer, and this membrane series is called XM series. This new membrane had smaller pore size than the previously used membranes in cell encapsulation and provides more protection to encapsulated cells as compared to the previously used macroporous membranes. Therefore, these newly developed membranes allowed the xenotransplantation without using immune-protecting medicines.

In earlier studies, XM-50 membranes were used in a bioreactor for culturing the cells. This application leads to the studies of using XM-50 hollow fibers as intravascular encapsulation devices where these hollow fibers are loaded with cells namely islets of Langerhans and placed into the abdominal cavity of rats [186], while in some other instances, XM-50 was also used as extracellular devices where the cells were placed into the lumen of the hollow fibers and then both ends of the fiber were sealed by adhesives [187]. In earlier studies, the main focus was the rats and pigs. However, in an application, XM-50 hollow fibers were used for the first time as extravascular device in which cell suspension from the brain tissues from various animals including human was infused into the fiber, and then it was implanted into the brain of the rats suffering from limited growth due to the deficiency of growth hormone. It was found that the level of the growth hormone started to increase leading to the restoration of growth for 3 months [60]. These were the earlier developments on which the present era of macroencapsulation is based.

3.1.3 Cell Loading Chamber

The easiest way of loading a cell suspension in the encapsulation device is injecting the cell suspension directly into the device. However, sometimes this simple injecting technique results into the formation of cell clumps or aggregates as in case of aqueous suspension the driving force making the cells to settle down by gravity, which limits the uniformity in packing of large number of devices. Various efforts have been made to resolve this problem and to maintain the uniform cell suspension inside the devices. The removal of air becomes significantly important when the devices have only one opening. In this case, the air could remove by attaching a vacuum pump to the device or by introducing the venting mechanism in the device. In this process, the removal of the hydrophobic gas from the hydrophobic membrane is facilitated by wetting the membrane with a solvent, methanol. The wetting solvent is then removed by means of aqueous buffer maintained at the physiological pH. Following all these steps, the membrane becomes translucent and loading of cells could be viewed by means of stereomicroscope as the color of the membrane changes with loading of cells.

Another factor responsible for the loading of cells is the volume of encapsulation chamber. In some cases, large volume chamber is required because the amount of compounds released from the cells is usually very low; therefore, a large number of cells are required for chronic liver failure. But in some other cases the cells are required in smaller number because the cells can release the potent chemical agents for site specific action in the body [188]. Generally, the planar devices can hold the volume up to several millimeters, but in hollow devices for the encapsulation chamber can hold approximately 0.5–50 μL. Till to date, there is no method known that could exactly measure the encapsulation volume in the chamber making it difficult to predict the amount of biomass the hollow devices could retain.

3.1.4 Immobilization Matrices

One of the major problems of encapsulated cells is the necrosis of the cells (cell death due to tissue damage). The cells are dead with the passage of time and the amount of the pharmacological component starts to dwindle with time. To avoid this problem, the cells are loaded in the form of dilute suspension, but due to gravitational force acting on the cells the aggregates are formed slowly and exceed the transport limit of the membrane. This problem is resolved by immobilizing matrices that result into uniform dispersion of the cells in encapsulation chamber. For example, collagen containing gels along with glass beads is used in uniform spreading of the cells that require some support in the form of anchorage. On the other hand, hydrogels producing from alginate, chitosan, and collagen in which cells are suspended in the gels without an anchoring support. Both cationic and anionic matrices are used for these purposes. The neutral matrices such as highly cross-linked poly(ethylene oxide) and poly(vinyl alcohol) can also be used in cell encapsulation. These matrices help in uniform dispersion of cell suspension and prevent the formation of aggregates in the devices [189]. It is noteworthy to say that the viscosity of the matrix must be sufficient enough to allow the cells to disperse and eliminate the effect of the gravity avoiding the formation of the aggregates. Moreover, the matrix should also be mixed with such components that could provide food and nourishment to the cells.

3.1.5 Cell Engineering for Maximal Effectiveness

The cells used in transplantation could have different purposes in the host. For example, the cells may be used to release some hormone, a neurotransmitter for electrical signals, cytokines, growth factors, and growth inhibitors, etc. In most of the instances quickly proliferating cells, stem cells are used. Now a days, it is also possible to genetically engineer an encapsulated cell for the production of a specific compound required in the body of the patient. It is more feasible to genetically engineer the rapidly proliferating cells as compared to fully matured cells [190]. Another extremely important aspect of the encapsulated cells is to maintain the constant level of living cells after encapsulation because the encapsulated cells undergo proliferation and ultimate cell differentiation. Different views have been given to handle the issue of the cells, e.g., the number of cells could be maintained at constant level by using extracellular matrix to cause the cells to differentiate and prevent the proliferation of the cells. Latest approach has suggested the use of cell cycle regulators to control the proliferation of the cells so that they may not exceed the limit of the encapsulation membrane [191].

3.1.6 Long-Term Utility of Encapsulated Device

Demand of the medical sciences is the life of the encapsulated devices and the cells inside them. In many cases, the transplanted cells are required to stay in the body of the host for long time even for years [192]. In this case, there may be an instance where the device is required to remove from the host, and/or it may have to be replaced by another one. In order to fulfill this requirement, different approaches have been developed. The membrane should be mechanically strong enough to tolerate the force applied during its removal from the body. In one process, a tether is attached to the device and then it is transplanted, later on when there is need of removing the device from the body it could be pulled out. In order to make this device mechanically strong, a strain relief element is axially oriented in the fiber to avoid its breakage during eradication [193].

Another way is to replace the cells in the device instead of removing the entire device is to make an opening in the existing device replacing and/or refreshing the cells inside the device. In this case a suspension of cells is used. There is another approach to address the issue of retrieving the device. In this approach, the cells are mixed with polymeric solution that undergoes gelation just by changing the environment of the device. The polymeric solution undergoes gelation when the cells are transplanted into the host and hence the cells also become immobilized by the matrices formed by the polymeric materials [194]. Now it is possible to place, replace, or remove the cells from the interior of the devices without removing the device from the host.

3.2 Microencapsulation

Microencapsulation is a process in which tiny particles or droplets are surrounded by a coating to give small capsules of many useful properties. Different polymeric materials could be used for coating, e.g., polymeric matrices [195]. It could be seen that the cells are encapsulated by semipermeable membrane and matrix (Fig. 10). These cells are immobilized and supported by matrix which is also semipermeable and allows the exchange of essential delivery of the products for a longer period of time as cells release the products continuously (Fig. 1). Finally, the immobilized cells are surrounded by hollow device made up of semipermeable membrane that could be used for numerous applications. The encapsulated cells, called artificial cells, remain bound in the encapsulation membrane, which is permeable to vital components for the cells nutrients such as oxygen, nutrients, and electrolytes. While immunoactive components (antibodies, immunoglobulin, etc.) larger in size ranging from 160 to 900 kDa cannot enter into the capsule, the cells remain protected from the host immune system.
Fig. 10

Typical microcapsule showing the homeostasis of microencapsulated capsule [190]. (With kind permission of Wiley)

The eukaryotic cells have been implanted for more than 50 years since the first erythrocytes were encapsulated in 1966 by nylon microspheres [196]. The first application of microencapsulated cells was reported in 1980s when the encapsulated islets of Langerhans were implanted into the rat to treat diabetic. After a while, the first application in the human being was found in 1994 when the islets of Langerhans were implanted in a diabetic patient to release insulin. It was found that the glucose level was maintained in the patient for 9 months [197].

Cell microencapsulation is a technology with enormous clinical potential for the treatment of a wide range of diseases. In cell encapsulation system, cells are usually surrounded by liquid. Microcapsule can be prepared in different ways, which could potentially be used in cell encapsulation.

3.2.1 Microencapsulation by Polyelectrolyte Complexation

In this technique, the cationic and anionic polymers are used to form membranes for encapsulating the cells. If the polymeric compounds are soluble in the aqueous medium, then it becomes very convenient to carry out the encapsulation of the living cells. Both natural and synthetic polymers are used for this purpose. Natural polymers are considered to be the best polymer for encapsulation as they are more biocompatible to the cells. Among the natural polymers, alginate is the best studied and highly suitable material for encapsulating the cells. However, alginate has a severe limitation, i.e., the isolation of the alginate from the living systems at a massive scale gives highly heterogeneous composition. Both of these factors contribute to the viscosity of solution affecting the production of microcapsules [198]. Nevertheless, these days a number of techniques have been developed for harvesting the homogeneous samples from the natural sources. Synthetic polymers on other hand could be produced on a large scale with uniform composition, but it has a problem of biocompatibility for the encapsulation of the cells.

One of the most widely studied microencapsulation system using polyelectrolyte complexation involves alginate-poly-l-lysine-alginate (APA) microcapsules derived from the protocol of Lim and Sun [199]. The alginate-PLL system is developed by complexation reaction between polyanionic alginate and polycationic PLL. Actually alginate-PLL system has been used extensively for encapsulating the islets of Langerhans, which has been used for treating the patients of Insulin Dependent Diabetes Mellitus (IDDM), Type-1 diabetes. Alginate is a linear binary copolymer composed of β-d-mannuronic acid (M) and α-l-guluronic acid (G) residues. In this technique, the cell containing alginate polyanions was extruded through a needle with concentric air flow dropping into the gelation bath which contains a cross-linking agent solution of calcium chloride. Calcium ionically cross link with guluronic acid residue of alginate and form Ca alginate beads (Fig. 11). The Ca alginate bead is then suspended in PLL solution resulting in the formation of selectively permeable Alg-PLL membrane. The thickness and porosity of the membrane depends mainly upon concentration and molecular weight of PLL and the time of contact between the PLL and alginate. Finally, the Alg-PLL microcapsules are further dipped in a solution of negatively charged species most likely alginate solution so that any remaining positive sites on PLL could be neutralized and this in turn not only makes the microcapsules durable but also biocompatible [200]. The islets encapsulated by alginate-PLL system are reported to have an important role in normoglycemia in rodents and humans. The duration of the effectiveness of encapsulated islets varies from days to months and even for years in rodent [201]. However, these microcapsules show insufficient strength when implanted into larger animals such as dogs and presumably humans [202]. This may be due to their high water content or the loss of the polyelectrolyte overcoats.
Fig. 11

Schematic representation of entrapment of cells in gel beads [13]. (With kind permission of SciTech)

3.2.2 Encapsulation Based on Interfacial Phase Inversion

A new technique has been developed to encapsulate the cells using a thermoplastic polymer with desirable properties when transplanted into the host. The main concern with these thermoplastic polymers using in cell encapsulation is the insolubility in water. In such case, the spraying method can be used for producing matrix type cell encapsulated microcapsules [203]. During free falling of cells-polymer suspension droplet that generated by blowing the air, a rapid interfacial reaction can occur by spraying an aerosol cross-linking solution towards it, which gives rise to liquid shell (Fig. 12). Droplets containing cells are generated by blowing the air, but this may exert shear force on the droplet causing the variation in size and damage to the individual droplets. The viscosity and density of the liquid in the phases is adjusted to maintain the desired uniformity of the droplets, approximately 400 μm [204].
Fig. 12

Schematic representation of encapsulation of cells using the spraying method [13]. (With kind permission of SciTech)

The advancement in the interfacial cell encapsulation has led to the application of a new polymer namely hydroxyethyl methacrylate-methyl methacrylate (HEMA-MMA) in a ratio of 75:25. This polymer has led to the development of desirable characteristics such as permeability and mechanical strength of the microcapsules. This polymer has also shown promising results in dopamine released from the encapsulated cells that cure cerebral diseases [205].

3.2.3 Encapsulation by In Situ Polymerization

This technique was developed by Dupuy for encapsulating the cells in a membrane [206]. In this technique, a cell is surrounded by a healing polymer, which turn into a microcapsule with moderate strength. When these cells are transplanted, the microcapsule can be damaged and eventually ruptured which then triggers the healing polymer to undergo further polymerization to keep the cell protected. This technique provides the cells with much longer shelf life than ordinarily encapsulated cells. In recent past, a system has been developed in which microcapsules are made up of urea-formaldehyde (UF) to surround the healing agent, dicyclopentadiene (DCPD) [207].

3.2.4 Conformal Coating Techniques

This technique is quite advantageous as compared to aforementioned cell encapsulation approaches. It has been observed in encapsulated islets that show limited release of insulin in response to glucose when implanted in the body. The main reason behind this insufficient release of insulin is the limited diffusibility of molecules across thick and large sized (600–1000 μm in diameter) capsules. Additionally, the exact implantation of the islets into the target area of the body is also affected by the large sized molecules. In order to address these issues, a new technique called conformal coating has been developed to minimize the size of the capsule as well as that of grafting tissues to transplant them into the desired location.

The coatings are considered conformal as they fit to the size and shape of each cell. This is especially needed in case of islets because these cells vary in their shapes and sizes. In this technique, PEG hydrogel precursor is directly applied to islets to adjust according to their geometrical aspects. After PEG has surrounded the islets, the precursor is cross-linked into an elastic hydrogel. The cells are first surrounded by small amount of the coating precursor which is then cross-linked to form quite strong gel, also called shrink wrapping. In this way the cells get encapsulated in a thin layer of capsule that would give stability and uniformity to the islets making them more applicable [208].

3.3 Other Cell Encapsulation Techniques

3.3.1 Silica Sol-Gel Entrapment

Sol-gel chemistry is used to develop a new type of membrane based on ceramics. These ceramic-based membranes are completely novel in the cell encapsulation technology. The significance of these membranes lies in the ease of synthesis of these membranes from a solution of raw materials at room temperature and normal pH of the living cell. Secondly, their pore size could also be controlled easily even when the membrane is applied to the living cell. Additionally, the cells could maintain their normal physiological activities even under varied ionic concentrations, and cells could survive for 6 months in the presence of trypsin in an in vitro analysis [209]. The membranes spheres are synthesized by drop-tower sphere generation and emulsion polymerization. The pore size of the membrane was controlled in such a way so that only insulin and other biological molecules like cytokines can pass through while antibodies are not allowed to pass through these membranes [210]. Furthermore, the cells encapsulated in a ceramic membrane have good sustainability when transplanted in vivo and their recovery from the host has also shown that they were active even after 1 month. Although the cells encapsulated by sol-gel entrapment technique have not shown complete normoglycemia in some cases, they are potential alternative for other methods for different applications such as drug delivery and treatment of diabetes. They could also provide mechanical strength and durability to the cells in the host [211].

3.3.2 Thermo-reversible Gelation

This technique is based on the gels prepared by changing the temperature of the polymer solution. The aqueous solution of thermo responsive polymeric gel converted into a solution when the temperature of the polymer solution is lowered. However, when temperature is raised, these polymers form a coil in the solution which in turn changes into a three dimensional network and appears as a gel. However, there are some exceptions to this common behavior of naturally occurring thermo responsive polymers, i.e., they are liquid at room temperature but turns into a gel on heating [212].

This sol-gel transformation is utilized for developing gels that easily encapsulate the living cells. These gels especially made from methylcellulose could have shown considerable effect on the repair of the brain cells [213]. In another application, the cells encapsulated by this technique have shown sustained drug release for 60 min. But sometimes surfactants were used to stabilize the microcapsules, which could result into toxicity, and limit their applicability in an in vivo application [214].

3.3.3 Chondrocyte Membrane

This technique is based on a completely novel concept, i.e., instead of encapsulating the cells in a physical barrier, biological mechanism is used to protect the cells from the immune system. In general, the immune privileged cells, which are not attacked by the immune system of the body, for example, chondrocytes, are used to prepare the membrane [215]. In this process, the islets are entrapped within the capsule made up of chondrocytes. In a procedure developed by Pollok et al., islets were seeded on biodegradable polyglycolic acid polymer [216]. Afterwards the cells were encapsulated with monolayer of chondrocytes. The encapsulated cells along with controls were kept in culture for 5 weeks in which one group of cells were fed with glucose every 5 days. The amount of insulin in culture medium was monitored. It was found that islets were functional. This indicates that chondrocyte encapsulation membrane permits diffusion of glucose and insulin with intact structure of the islets [217].

As a whole, the debate between macro- versus microencapsulation is an ongoing dispute, and neither technique has demonstrated clear superiority over the other. Above all, the success of any encapsulation technique ultimately relies on a systematic evaluation of capsule properties and encapsulated cell performance.

4 Assessment of Capsule Properties

The capsular properties are one of the many important factors that were accessed by the interaction properties of the capsule and the performance of encapsulated cells. Understanding the relationship between the performance of the encapsulated cells and the encapsulating capsules is vital in the development of a successful technology for cell encapsulation. In spite of the achievement of the desired aim of cell encapsulation, in depth understanding of the capsules could result into much improved properties such as permeability, mechanical properties, immune protection, and biocompatibility of encapsulated capsules.

4.1 Permeability

There are two factors of prime importance that determine the survival of the living cells. The cells permeability determine which compounds will enter and exclude by the cell, and the molecular weight cut-off (MWCO) will define the upper limit of the molecules getting into or out of the cells. The major hindrance is the inability to measure the actual pore size of the capsule. The indirect methods have been used to determine the pore size of the capsules. The permeation rate of different solutes into the cells can give estimation about the pore size of the capsule [218]. Some successes have been achieved in measuring the size of the macrocapsules; however, much efforts are being needed to determine the pore size of the microcapsules accurately.

4.1.1 Microcapsules Permeability

The pore size of the capsule could be determined by simply equilibrating the capsule with the known volume of desired solute and then the time taken by the solute to get into the cell and to reach equilibrium is measured [219]. This gives mass transfer of the solute which is used to determine the capsule permeability in the following manner:

The relationship between the rate of mass transfer, U, and the membrane permeability, Pm is given by [220]:
$$ {\mathrm{P}}_{\mathrm{m}}=\mathrm{U}\times \mathrm{d} $$
where, Pm = permeability of the capsule; U = rate of mass transfer as function of time; d = thickness of the capsule membrane.

The abovementioned technique gives the permeability of the capsule and provides some qualitative information about the pore size of the capsules. Another technique has recently been developed called inverse size exclusion chromatography to quantify the pore size of the microcapsules [221]. In this technique, the column is packed with microcapsules and solute is loaded into the column to pass over the capsules with the loaded mobile phase, and partition coefficient, Ksec, is measured which indicates the pore size distribution in the capsule membrane.

In another procedure for determining MWCO, 1 mL of the microcapsules is mixed with 1 mL of 0.05% (w/v) commercial dextran-FITC samples with nominal MWs of 10, 70, 150, 250 and 500 kDa. The microcapsules are then incubated for 24 h at 37 °C and observed with a confocal laser scanning microscope (CLSM). The intensity profile obtained by CLSM showed increasing in-diffusion with decreasing MW (Fig. 13), and the MWCO of the microcapsules were found to be 100–200 kDa [222].
Fig. 13

Top: CLSM middle sections of cell containing capsules exposed for 24 h at room temperature to 0.05% dextran-FITC with nominal MWs of (a) 10 k, (b) 70 k, (c) 150 k, (d) 250 k, and (e) 500 k. Bottom: Line profile from images as above [222]. (With kind permission of ACS)

4.1.2 Macrocapsules Permeability

The macrocapsule membranes are different from the microcapsule. The macrocapsule membranes are thicker and have the ability to withstand high pressure. The properties of these membranes depend upon the porosity and the ability of the membrane to retain a molecule of a particular size in the membrane is expressed in the form of nominal MWCO (nMWCO). The MWCO is the MW of the largest molecule that is permitted to pass through the membrane, whereas the nMWCO indicate that the 90% of the molecules of particular size are not allowed to pass through the membrane under higher pressures [223].

In general, the transport properties of the macrocapsule membranes are determined by hydraulic permeability assessment, a technique based on consumption of ultrapure water. In this technique, the testing device consists of a flow rate controllable pump, membrane holder, pressure transducers, and balance that measure flow rate. Generally a lower pressure is maintained to reduce the chances of membrane deformation. The hydraulic permeability measured in this process is reported as unit of flow per unit surface area per pressure unit. This would indicate the pore size of the macrocapsule membrane [224].

4.2 Mechanical Properties

The determination of the mechanical properties of the capsules is important for the production. These properties usually give an indication about the strength as well as the integrity of the capsule. The durability of microcapsules is usually tested by subjecting the capsule to shear flow. The number of “failed” capsules is simply a measure of mechanical durability of the capsule membrane. The mechanical force faced by the capsule depends on the shear rate and viscosity of the fluid in this system. No doubt the method is very simple yet it is used as a screening tool to improve the process of capsule manufacture and produces capsules with better mechanical properties such as coating time, surface modification, and choice of polymeric additives to capsule membrane [225].

The mechanical properties of the capsules are generally determined by compression test [226]. In dry test, capsules are drawn into a pipette causing the release of a single capsule on the surface of the plate. An image of the capsule is taken before the compression test for determination of the initial diameter of the capsule, then the capsule is introduced to the apparatus for the compression test. In the compression machine, the capsule and punch is kept separated until the stepper achieves a steady state velocity. The compression program is started after positioning the punch above the capsule, and it gets terminated after failure of the capsule is observed. The immersion test is also performed modifying the dry test [227]. In case of immersion test, the capsules are dispersed in a bath of dicyclopentadiene (DCPD) and allowed to equilibrate for 24 h. As the capsule is dipped for long time in the bath, some diffusive mobility across the capsule membrane is observed. Afterwards a single capsule is taken by pipet and placed into the compression cell and small quantity of the fluid is added to ensure that capsule is completely immersed. The rest of the procedure is exactly the same as done for dry test to obtain the mechanical strength of the microcapsules [228].

4.3 Immune Protection

The immune protection of the transplanted cells depends on the diffusion of antibodies into the cells or onto the surface of the cells [229]. The diffusive permeability of the encapsulating membrane is dependent on the donor-recipient mismatch, i.e., allogeneic versus xenogeneic tissues. In case of allogeneic graft tissue, reactive immunoglobulins are absent and hence the host’s immune reactive cells directly reach the allogeneic graft tissue which need to be prevented by the encapsulating membrane. In xenogeneic tissue, there are antigens derived from their surface and hence they activate host’s immune system especially CD-4 cells which in turn produces antibodies that stick to the grafted cell’s surface and interferes with the diffusion properties of the membrane [230]. Therefore, the approach for determining the diffusivity in allogeneic graft cells is called the direct pathway, whereas in case of xenogeneic graft tissue, the approach is called indirect pathway and is of secondary importance for allogeneic graft cells.

In actual practice, it becomes almost impossible to avoid the activation of host’s immune system because of the failure of small portion of the grafted capsules [231]. In this way, the direct pathway can be activated which could attack the failed grafted cells and hence the number of graft specific T-cells increases in the grafted area. The increase in the number of graft specific T-cells could be measured as a parameter for determining the cell’s diffusivity, but it is not clear up to what level these T-cells will affect the intact capsules in the body [232]. Another concern in direct approach is the occurrence of cytotoxic agents in the grafted tissue area. The cytotoxicity of the grafted cells could be measured by in vitro cell culture techniques. However, if the direct pathway is activated as an inadequacy in the capsule properties, then conditions for determining the diffusive properties of the cells will become highly stringent.

The permeability of the antibodies is simply measured by enhancing the transfer of the antibodies into the cells which is achieved by making an antibody sink in the membrane, but extra care should be taken if the large sized proteins are present in the antibody mixture. However, this issue could be resolved by using high purity proteins in the analysis and encapsulating membranes that are very reactive towards antibodies. The amount of the antibodies that have reacted with the capsule gives an indication about the response of the antibodies to the cell. Another way to determine immune protection of the capsule membrane is to mix the capsules with lymphocytes [229, 233]. However, the success of the techniques lies in the analysis of the immune protection in in vivo experiments as the cells are to play their final role in the body of the host. Nevertheless, the correlation between the in vitro and in vivo experiments could be very useful in determining the possible protection of the encapsulated cells in the body of the organism.

4.4 Biocompatibility Assessment

Biocompatibility of any biomaterial is defined by its performance in the body of the host with a tolerable host response in a particular application, e.g., the transplantation of artificial organs such as artificial breasts and hips, etc. [234]. A biomaterial is considered fully biocompatible if the cellular system having encapsulating membranes produces no or not more than a minimal foreign body reaction [235]. The response of the host is a serious problem to the clinical application in an encapsulation technology. One of the serious consequences of nonbiocompatibility of encapsulated cells is the use of fibrotic overgrowth on the encapsulated cell’s surface affecting the transport of the molecules in and out of the encapsulated tissue [236]. The other factors that play a role in making the encapsulated cells biocompatible are the methods of surgical implantation, size of the device, morphology of the cell surface, the material composing the macroencapsulation device, and the content shed from the encapsulation chamber [200, 237]. When the device is implanted, the host immune response is initiated as a result of disruption of host’s vasculature. Initially, the cell surface is covered with activated platelets, humoral serum components, clot constituents, cell debris, and extracellular matrix. All the debris from the cell surface is removed by the tissue macrophages and the process of wound healing is initiated. In the final step, the mesenchyme cells produce the matrix and neovascularization around the implant completes the process [238].

The production of cytotoxic agents from the host cells also plays a role in response to the grafted encapsulated cells. The material of the capsule membrane plays a significant role in determining the response of the host [239]. Several studies have been conducted on these toxic species, and it has been found that there are two types of toxins mainly one with short life spans such as oxygen free radicals and second one are more stable and can permeate through the membrane of the capsule, e.g., cytokines [240]. Recently, several studies have shown material dependent production of cytokines in encapsulated cells [241]. The assessment of cytokine production can be carried out by using macrophages from peritoneal cavity, the site that is mostly utilized for implantation of the encapsulated cells [242].

Another significant aspect in encapsulation implants is the neovascularization around the grafted tissue in the host. It has been found that the outer topography of the membrane plays major role in determining the neovascularization response [243]. Generally, the membranes that have larger surface pores result in neovascularization very close to the host-material interface. The formation of a huge number of blood vessels is not desirable in many clinical applications as they may hinder the release of the desired therapeutic agent from the encapsulated cells [244]. In some applications, there might be a desire to maintain a high concentration of therapeutic agents systemically. In other applications, it is quite possible to implant the cells in the body for a shorter period of time and subsequent retrieval after the correction/cure the problem. In this situation, presence of huge number of blood vessels will hinder in the retrieval of encapsulated devices from the body of the host [245].

5 Efficacy and Survival of Encapsulated Cells

It has been found that a slight variation in the procedure of encapsulation could result into an important challenge in the performance of encapsulated cells. There is one common problem associated with encapsulated islets is that the implanted islets have never survived permanently into the host body [183]. This decrease in the efficacy of the implanted tissue is attributed to overloading amount of glucose in the blood and the less quantity of the implanted tissue used in the host. Even after using large number of encapsulated cell, De Vos et al., showed that there was decrease in the efficacy of the implanted tissue with the passage of time, which indicates that the efficacy of the transplant cells neither depends upon the load of sugar nor the insufficient number of islets, and there are some other factors also responsible for deficiency in the efficacy of the implant [246].

A formulated hypothesis states that the number of renewing islets is not equal to the number of necrotic islets. Later this hypothesis was experimentally tested, and it was found that the actual reason is the life span of the insulin producing β-cells. These β-cells were not able to replicate after 3 months, and there were decreased in the number of β-cells leading to decrease in the efficacy of the implanted islets. The cell death is found to be more in the center of the implant than at the periphery. The reason for this behavior is the inability of the nutrients to arrive at the center of the implant [247].

6 Challenges in Cell Encapsulation

In spite of the abovementioned successes, there are still many challenges and promises that cell encapsulation has to meet for touching the new horizons in the medical sciences. For the last few decades, the researchers have been working day and night on developing the systems that have the ability to release the drug in a highly controlled manner so that the effectiveness of the drug could be improved with minimal side effects leading to the improved quality of life. Among these techniques, cell immobilization is the best approach whereby the cells entrapped and immobilized in the matrices act as the factories of therapeutic agents curing various ailments. The promising potential of cell encapsulation lies in the fact that the cells are not only made into drug releasing factories but also eliminate the need of administering immune-protective drugs during transplant of cells.

The history of cell encapsulation is full of many successes and defeats. At one edge, the continuous release of therapeutic agents has been proven to be successful in curing numerous diseases in animal models such as the diseases due to hormonal deficiencies, hemophilia, diseases of CNS, and cancer. Additionally, the technique has also been reported in many clinical trials for applications in curing human diseases. However, the general feeling is that this technique is not meeting up to the expectations. No doubt scientists from across the world have focused much of their efforts on the field, but the reality is that there is no product in the market until now due to the nonreproducibility of results in animal models, lack of standardized technology, immediate necessity of reproducible, and biocompatible materials for making biocompatible and stable devices. Moreover, their systematic administration is associated with unwanted side effects due to nonspecific suppression of the immune system that leads to a variety of undesired complications (e.g., opportunistic infections, failure of tumor surveillance, as well as adverse effects on the encapsulated tissues). Furthermore, the controlled release of drugs by the encapsulated cells is also another important consideration if the promises of cell encapsulation are to be met. These ditches happening in the field of cell encapsulation have changed the attitude of many scientists leading to stepwise understanding of encapsulation technology. This stepwise approach is important to answer the research questions leading to in depth understanding of the factors that are responsible for limiting the success and help renewing the excitement and hopes surrounding this cell-based technology.

7 Conclusions

Nowadays the encapsulation technique is touching new horizons where encapsulated cells are playing role in gene therapy and drug delivery system in the human body. There are a large number of technologies that have been developed over the past years to improve the techniques and their outcomes in an effective manner. Cell immobilization and encapsulation has a wide range of applications. It appears likely that by the end of the decade clinical trials of encapsulated cells to treat many of these diseases will become a reality. It is worthwhile to mention that the biocompatible materials with much improved properties could serve the desired purpose in cell encapsulation. A better and in depth understanding of the host immune response could also result in developing appreciable biomaterials for cell encapsulation. A much better comprehension of the properties of the biomaterials with much control on the properties of the biomaterials during fabrication could lead to clinical applications of cell encapsulation. However, the biotechnological development including the sourcing of raw materials, the design and building of manufacturing facilities, the scale-up and optimization process, storage and distribution of the product, and quality control thus far is time-consuming, which requires collaboration between scientists and engineers from many disciplines. A discussion about the future application of encapsulated cells, design systematic protocol, and important objectives for the scientists and industries is yet to be considered for the benefit of all.

Notes

Acknowledgments

The authors would like to gratefully acknowledge King Fahd University of Petroleum & Minerals (KFUPM), Saudi Arabia for providing excellent research facilities.

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Copyright information

© Springer International Publishing AG, part of Springer Nature 2018

Authors and Affiliations

  • Abdul Waheed
    • 1
  • Mohammad Abu Jafar Mazumder
    • 1
    Email author
  • Amir Al-Ahmed
    • 2
  • Partha Roy
    • 3
  • Nisar Ullah
    • 1
  1. 1.Chemistry DepartmentKing Fahd University of Petroleum & MineralsDhahranSaudi Arabia
  2. 2.Center of Research Excellence in Renewable EnergyKing Fahd University of Petroleum & MineralsDhahranSaudi Arabia
  3. 3.Department of Pharmaceutical TechnologyAdamas UniversityKolkataIndia

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