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Keywords:

  • therapeutic protein;
  • cell viability;
  • biocompatibility;
  • biomolecular transport;
  • stiffness

Abstract

  1. Top of page
  2. Abstract
  3. INTRODUCTION
  4. CELL ENCAPSULATION PROCESS
  5. CONTROL OF HYDROGEL IMMUNOGENICITY
  6. CONTROL OF TRANSPORT PROPERTIES IN HYDROGELS
  7. CONTROL OF MECHANICAL PROPERTIES OF HYDROGELS
  8. CONCLUSION AND FUTURE DIRECTION
  9. References

Stem cells, progenitor cells, and lineage-committed cells are being considered as a new generation of drug depots for the sustained release of therapeutic biomolecules. Hydrogels are often used in conjunction with the therapeutic secreting cells to provide a physical barrier to protect the cells from hostile extrinsic factors. Although the hydrogels significantly improve the therapeutic efficacy of transplanted cells, there have been no successful products commercialized based on these technologies. Recently, biomaterials are increasingly designed to provide cells with both a physical barrier and an extracellular matrix to further improve the secretion of therapeutic proteins from cells. This review will discuss (1) the cell encapsulation process, (2) the immunogenicity of the encapsulating hydrogel, (3) the transport properties of the hydrogel, (4) the hydrogel mechanical properties, and will propose new strategies to improve the hydrogel and cell interaction for successful cell-based drug delivery strategies. © 2008 Wiley Periodicals, Inc. J Biomed Mater Res 2008


INTRODUCTION

  1. Top of page
  2. Abstract
  3. INTRODUCTION
  4. CELL ENCAPSULATION PROCESS
  5. CONTROL OF HYDROGEL IMMUNOGENICITY
  6. CONTROL OF TRANSPORT PROPERTIES IN HYDROGELS
  7. CONTROL OF MECHANICAL PROPERTIES OF HYDROGELS
  8. CONCLUSION AND FUTURE DIRECTION
  9. References

Over the past few decades, significant advancements in molecular and cell biology have enabled scientists to identify a number of chronic and malignant disease mechanisms and develop various therapeutic drugs. These therapeutic drugs, including small molecules, hormones, proteins, and genes aim to regulate specific cellular activities and lead toward the recovery of homeostasis. Unfortunately, drug therapies are often plagued by a rapid loss of bioactivity and subsequently limited therapeutic efficacies.1, 2 Various organic and inorganic drug delivery devices have been designed to deliver drug molecules in a sustained manner while minimizing drug denaturation, but these drug delivery devices have experienced limited successes.3

Recently, cells have been increasingly exploited as alternative drug delivery devices. Cells can act as drug depots enabling the delivery of therapeutic molecules over an extended time period.1 Cells are capable of delivering drugs in response to an external stimulus, which is highly advantageous to maintain homeostasis for patients suffering from chronic diseases, like diabetes or Parkinson's disease.4, 5 In addition, cells secrete therapeutic proteins and cytokines that may not be purified or synthesized in a test tube. This provides the potential to treat various diseases that cannot be cured with currently available technologies.6 A variety of stem cells, progenitor cells, lineage-committed cells, and genetically engineered cells are being tested in preclinical and clinical trials as drug delivery vehicles, as shown in Table I. Stem and progenitor cells secrete a diverse array of growth factors, including vascular endothelial growth factor (VEGF) and nerve growth factor, which are used to treat ischemia and neuronal damage, respectively.7, 8 Islet cells, kidney cells, and parathyroid cells are examples of lineage-committed cells, which naturally secrete insulin, erythropoietin, and parathyroid hormone, respectively. These therapeutic proteins are used in the treatment of diabetes, anemia, and hyperthyroidism.4, 9, 10 Many cells, such as kidney cells, ovary cells, fibroblasts, and myoblasts, have been genetically engineered to secrete specific therapeutic proteins. Such proteins include endostatin to suppress cancer metastasis, erythropoietin to treat ischemia, nerve growth factor to treat Alzheimer's disease, and β-endorphin to alleviate pain.11–15

Table I. Examples of Cell-Based Drug Delivery Tested in Preclinical and Clinical Settings
Cell-Secreted ProteinCell TypeTarget DiseaseType of TrialRef.
  1. VEGF, vascular endothelial growth factor.

InsulinIslet cellsDiabetesPreclinical4
DopaminePheochromocytoma and adrenal chromaffin cellsParkinson's diseasePreclinical5
VEGFHematopoietic stem cellsIschemiaPreclinical7
Neurotrophic factorsNeural stem cellsNerve damagePreclinical8
ErythropoietinKidney cellsAnemiaPreclinical9
Parathyroid hormoneParathyroid cellsHyperthyroidismClinical10
EndostatinEngineered human kidney cellsCancer metastasisPreclinical11
EndostatinEngineered hamster kidney cellsCancer metastasisPreclinical12
ErythropoietinEngineered myoblast cellsAnemia and ischemiaPreclinical13
Nerve growth factorBaby hamster kidney cellsAlzheimer's diseasePreclinical14
Beta-endorphinEngineered mouse tumor cellsAlleviate painPreclinical15

In treating diseases, the strategy of cell transplantation is as critical for the success of cell-based drug delivery as the ability of cells to secrete the desired drug molecules. Cells directly injected into the body experienced less than desired therapeutic efficacies for many reasons, including immune rejection and rapid decrease in cell viability.1 The immune response is most prominent for allogenic or xenogenic cells, but the need for plentiful cell sources dictates the use of these cell types. Immunosuppression drugs may reduce the risk of immune rejection of the transplanted cells, but as a side effect the body is more vulnerable to other infections.16 Furthermore, transplanted cells not properly protected from external mechanical loadings or not surrounded by the proper microenvironment may rapidly lose their viability and functionality to secrete drug molecules.17

Encapsulating cells into biomaterials has shown promising results for reducing the immune response and increasing the efficacy and viability of transplanted cells.4, 18 Biomaterials are commonly processed into hydrogels acting as permeable membranes to block the access of immune cells to the transplanted cells.19 These hydrogels also function as structural supports for the cells, protecting the cells against external mechanical loading. In addition, hydrogels are structurally similar to extracellular matrices and provided many advantages in encapsulating cells as compared with other biomaterial forms.20–22 Both naturally derived and synthetic polymers, including alginate, hyaluronic acid, agarose, poly(ethylene glycol) (PEG), and poly(hydroxyethyl)methacrylate (HEMA), are fabricated into cell-encapsulating devices.4, 23–26 Interestingly, these biomaterials are typically investigated for cell-based drug delivery with only a few different cell types, shown in Table II, despite the large variety of cell types which are able to secret proteins, as shown in Table I. Nevertheless, certain impressive results have been obtained from preclinical and clinical trials for implanting encapsulated cells, but many results are inconsistent because of the limited ability of hydrogels to support the cellular viability and function over extended time periods. Therefore, few successful products have been fully commercialized based on these cell-encapsulation technologies.

Table II. Examples of Biomaterials Used to Improve Cell-Based Drug Delivery
Purpose of BiomaterialBiomaterialCell-Secreted ProteinCell TypeTarget DiseaseType of TrialRef.
  1. PEG, polyethylene glycol; ESCs, embryonic stem cells; HEMA, 2-hydroxyethyl methacrylate; CNS, central nervous system disorders; CNTF, ciliary neurotrophic factor; VEGF, vascular endothelial growth factor; BMP, bone morphogenetic protein.

Improve cell viabilityPEGInsulinIslet cellsDiabetesPreclinical20, 25
Hyaluronic acidFibroblasts and ESCsPreclinical23
AlginatePreclinical27
AlginateOsteoblastsPreclinical28
InsulinIslet cellsDiabetesPreclinical29, 30
AlginateInsulinIslet cellsDiabetesPreclinical31
Decrease immunogenicityAlginateFibroblastsPreclinical32
HEMACNSPreclinical26
AlginateInsulinIslet cellsDiabetesPreclinical33, 34
PEGInsulinIslet cellsDiabetesPreclinical35–37
PEGInsulinIslet cellsDiabetesClinicalClinicaltrials.gov
PolyethersulfoneCNTFEngineered kidney cellsHuntington's diseaseClinical38
Control transportAlginateInsulinIslet cellsDiabetesPreclinical39, 40, 41
AlginateInsulinIslet cellsDiabetesClinical42
PEGInsulinIslet cellsDiabetesPreclinical43
Improve mechanical propertiesAlginateInsulinIslet cellsDiabetesPreclinical40
AlginatePreclinical44, 45–47
AlginateVEGFFibroblastsIschemiaPreclinical48
PEGBMPEngineered fibroblastsBone defectsPreclinical49

Thus, extensive efforts are being made to improve the performance of hydrogels and the subsequent therapeutic efficacy of cell-based drug delivery. This review article presents various strategies of hydrogel design in terms of the encapsulation process, material immunogenicity, porosity, and stiffness. The influence of these hydrogel properties on improving cell viability, cell function, and subsequent therapeutic efficacy, both in vitro and in vivo, as well as possible future strategies to improve the efficacy of cell-based drug delivery will be discussed.

CELL ENCAPSULATION PROCESS

  1. Top of page
  2. Abstract
  3. INTRODUCTION
  4. CELL ENCAPSULATION PROCESS
  5. CONTROL OF HYDROGEL IMMUNOGENICITY
  6. CONTROL OF TRANSPORT PROPERTIES IN HYDROGELS
  7. CONTROL OF MECHANICAL PROPERTIES OF HYDROGELS
  8. CONCLUSION AND FUTURE DIRECTION
  9. References

The cell encapsulation process is a major factor for successful cell-based drug delivery because it greatly affects cell viability. The cell encapsulation processes can be largely divided into two categories: (1) macroencapsulation to encapsulate massive amounts of cells into one large volume device, such as hollow fibers, and (2) microencapsulation to encapsulate cells into multiple small volume devices, such as microsized beads.

Hydrogels are commonly processed into spherical beads ranging in size from 100 to 2000 μm for the microencapsulation of cells. The spherical nature of the hydrogel beads maximizes the surface area, and the small volume of the gel beads facilitates biomolecular transport. The procedure of microencapsulation depends on the cross-linking mechanism to form the gel. Cells have been readily encapsulated in calcium cross-linked alginate beads by extruding droplets, consisting of cells and an alginate solution, into the cross-linking calcium chloride solutions.27 Microfabrication techniques have been used to encapsulate cells in hyaluronic acid or PEG hydrogel beads formed by radical cross-linking. One of these techniques consisted of a mixture of cells and uncross-linked polymer solutions loaded into micromolds and polymerized in situ.23 The effect of various cross-linking molecules on the cell viability remains an often researched topic.

The rheological properties of the solution, prior to gelling, are crucial to maintain cell viability and cell–cell adhesion during the encapsulation process. Cells are often mixed with polymer solutions prior to cross-linking the polymer. Mixing cells with highly viscous solutions leads to a significant decrease in cell viability, because the high shear stress can physically damage cell membranes. In addition, high shear stresses can disrupt cell–cell contacts. The loss of cell–cell contact has been shown to reduce the viability of encapsulated β-cell within PEG gels.19, 25 The problems associated with mixing cells and polymer solutions may be tempered by engineering the molecular structure of polymer chains to decouple the dependency between the viscosity of the uncross-linked polymer solution and the mechanical properties of the gelled polymer. For example, a strategy to decrease overall molecular weights of the polymer molecules, while maintaining the number of cross-linking sites, allowed the preparation of hydrogel beads presenting high cell viability.28

CONTROL OF HYDROGEL IMMUNOGENICITY

  1. Top of page
  2. Abstract
  3. INTRODUCTION
  4. CELL ENCAPSULATION PROCESS
  5. CONTROL OF HYDROGEL IMMUNOGENICITY
  6. CONTROL OF TRANSPORT PROPERTIES IN HYDROGELS
  7. CONTROL OF MECHANICAL PROPERTIES OF HYDROGELS
  8. CONCLUSION AND FUTURE DIRECTION
  9. References

The success of cell-based drug delivery using hydrogels greatly depends on minimizing the patients' immune response, because the immune response is detrimental to cell viability, hydrogel stability, and mass transport. Following the implantation of cell-encapsulating hydrogels, the immune response is activated by the adsorption of proteins onto the materials, which will subsequently stimulate the recruitment of immune cells, such as macrophages.32 These immune cells directly destroy the transplanted cells through one of the many well-documented pathways.50, 51 The immune response causes structural damage of certain hydrogels labile to several intrinsic factors (e.g., enzymes) and extrinsic factors (e.g., external mechanical loading).52, 53 The immune response also elicits fibrosis around the hydrogels, which subsequently starve any encapsulated cells and limit the efflux of bioactive molecules secreted from the cells.54 Thus, most approaches developed to minimize the immune response to hydrogels are focused on preventing protein adsorption and cellular adhesion to hydrogels through the (1) encapsulation of cells into biologically inert hydrogels, (2) modification of the hydrogel surface with biologically inert materials, (3) anchorage of biologically inert molecules to the cell membrane, and (4) control of the hydrogel microstructure.

The immune responses to cell-hydrogel constructs are mediated with the use of biocompatible materials, which present minimal amounts of toxins and prevent protein adsorption. For example, cell-encapsulating calcium cross-linked alginate hydrogels have been extensively tested in many preclinical and clinical trials for the past few decades, because alginate molecules are anionic polysaccharides and do not associate with many proteins.55 Inhibiting protein adsorption onto the hydrogels may hinder the migration of immune cells and subsequently attenuate the threat to the encapsulated cells.56 However, the immune response to alginate hydrogels in vivo is controversial, because residual chemicals used to isolate alginate molecules from its natural source act as toxins. A rigorous purification process reduces the toxin levels and decreases the immune response to alginate hydrogels.57, 58 Encapsulation of islet cells in the purified alginate hydrogels significantly improved the therapeutic efficacy of cells by maximizing the viability and minimizing the fibrosis in the surrounding tissue.59 Other biocompatible hydrogels for the encapsulation of cells may consist of PEG or modified HEMA polymers, which inhibit protein and cellular adhesion to the material.60, 61 The development of a biologically friendly cell encapsulation process for these hydrogels (e.g., microfabrication-based process) may expedite the use of these hydrogels for cell encapsulation.62

Biocompatible materials are often used to modify the surface properties of certain cell-encapsulating hydrogels and reduce the immune response. Certain hydrogels, which facilitate the cell encapsulation process and provide the desirable physical properties, may induce severe immune responses. Alginate molecules are often used to coat hydrogels containing polycations, such as poly-L-lysine and poly(ornithine). Poly-L-lysine and poly(ornithine) allow for efficient control of pore size and mechanical properties but induce mild to severe immune responses.63 Overlaying alginate molecules strongly associate with polycations to form polyelectrolyte complexes and decrease the immune response, as depicted in Figure 1. The polycationic surfaces are also modified by chemically coupling PEG molecules to poly-L-lysine. Like alginate molecules, PEG molecules sterically inhibited the protein adsorption and cell adhesion onto poly-L-lysine-presenting hydrogels in vivo.32, 33 These surface treatments improved the therapeutic efficacy of islet cells by reducing fibrosis around the cell-hydrogel constructs. Similar results are expected with materials modified with other biocompatible polymers such as hyaluronic acids, cellulose sulfate, poly(ethersulfone), and HEMA.

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Figure 1. Schematic description of the strategies to control hydrogel immunogenecity. (a) Cell-encapsulating biomaterials may allow immune cells and proteins to adhere. (b) The use of biocompatible materials minimizes the protein adsorption and cell adhesion. (c) Coating of biomaterials with biocompatible molecules also prevents the protein adsorption and cell adhesion. [Color figure can be viewed in the online issue, which is available at www.interscience.wiley.com.]

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The immune response is also mediated by coating individual cell membranes with biocompatible polymers. PEG chains have been polymerized on the cell membrane surface of islet cells by adsorbing radical initiators (e.g., eosin) to the cell membranes. The presence of the thin PEG layers on the cell membrane did not affect the cell's ability to secrete drug molecules while providing a thin biochemical and biophysical barrier to the immune system almost 3 months.35–37 Individually coated islets with thin PEG layers have shown excellent preclinical results, and the technology is currently in phase I/II clinical trials (www.clinicaltrials.gov). Combining this cell-membrane coating technique with cell encapsulation techniques using biocompatible polymers may further minimize the immune response against cell-carrying hydrogels and extend the lifetime of transplanted cells.

The biological inertness of the cell-encapsulating hydrogels should be controlled and integrated with the microstructure. Nanoporous hydrogels can effectively prevent the invasion of immune cells. However, cytokines and immunoglobulin produced by immune cells often infiltrate the biocompatible gel matrix and also stimulate the recruitment of other immune cells. Islet cells encapsulated in calcium cross-linked alginate gels and implanted in rabbits encountered a significant immune response only after 4 days. The infiltration of immune cells decreased the therapeutic efficacy against diabetes.34 Also, in a phase I clinical trial, kidney cells engineered to produce ciliary neurotrophic factors for the treatment of Huntington's disease and encapsulated in poly (ethersulfone) with a molecular weight cutoff (MWCO) of 280 kDa also encountered significant cell death rates.38 These severe immune responses might have been significantly reduced by more thoroughly removing the toxins associated with these polymers. Overall, hydrogels may not be able to protect encapsulated cells from the immune system unless the pore size excludes these molecules. Conversely, too small of a pore size may limit the influx of biological molecules crucial to cell viability and the efflux of therapeutic drugs secreted from cells, as will be discussed in the next section. These complications related to biomolecular transport, however, may be resolved by integrating thorough purification of gel-forming polymers with refined control of pore size.

CONTROL OF TRANSPORT PROPERTIES IN HYDROGELS

  1. Top of page
  2. Abstract
  3. INTRODUCTION
  4. CELL ENCAPSULATION PROCESS
  5. CONTROL OF HYDROGEL IMMUNOGENICITY
  6. CONTROL OF TRANSPORT PROPERTIES IN HYDROGELS
  7. CONTROL OF MECHANICAL PROPERTIES OF HYDROGELS
  8. CONCLUSION AND FUTURE DIRECTION
  9. References

The success of cell-encapsulating hydrogels is also determined by their abilities to facilitate the influx and efflux of biological molecules into and from the encapsulated cells, as mentioned in preceding section. Without careful modification of the hydrogel microstructure, hydrogels may act as a physical barrier to limit the influx of nutrients and oxygen toward the encapsulated cells and prevent the efflux of therapeutic molecules and cellular wastes away from the encapsulated cells. Typically, the biological transport through the hydrogels is regulated on the basis of pore size, because the diameter of the pores determines the MWCO of the biomaterial, as depicted in Figure 2.

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Figure 2. Schematic description of the strategies to control transport properties. Optimal diameters of the surface pores (a) and bulk pores (b) of cell-encapsulating biomaterials facilitate the biomolecular from and into encapsulated cells while minimizing the infiltration of immune system. In contrast, too small surface pores (a) and bulk pores (b) inhibits the transports of all essential biological molecules. [Color figure can be viewed in the online issue, which is available at www.interscience.wiley.com.]

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Extensive efforts are being made to control the size of the hydrogel surface pores to allow the transport of oxygen, nutrients, therapeutic drug molecules, and cellular wastes, while preventing the transport of immunogenic molecules and cells. Specifically, the diameter of surface pores for an calcium cross-linked alginate hydrogel commonly varies from 5 to 200 nm.53 This pore size prevents the diffusion of large molecules, like fibrinogen (MW of 3.41 × 105 g/mol), but allows the diffusion of small molecules, like albumin (MW of 6.9 × 104 g/mol).64 This pore size is further controlled by forming polyelectrolyte complex layers on the gel surface. Exposing alginate hydrogels to polycations such as poly-L-lysine reduces the diffusivity of small molecules, like albumin or hemoglobin (MW of 6.8 × 104 g/mol) up to 50%. The ability of poly-L-lysine to reduce the effective surface pore diameter is dependent on the molecular weight of poly-L-lysine and the exposure time. Decreasing the molecular weight of poly-L- lysine molecules and extending exposure time led to a significant decrease in the average diameter of pores on the gel surfaces.39 Coating alginate gels with poly-L-ornithine (PLO) further reduced the permeability of a 75-kDa protein 10 times over poly-L-lysine.40 These approaches may be broadly applicable to cell encapsulation with other anionic polyelectrolyte biomaterials. Specifically, hyaluronic acid is being used with poly-L-lysine for the synthesis of multilayered films on solid surfaces.65

However, the presence of polycations on the surface of hydrogels-encapsulating islet cells induced protein adsorption and subsequently decreased the therapeutic efficacy of the cells.66 Gradual loss of cell viability due to the immune response was often observed from in vivo experiments.67 Specifically, PLO-coated gels elicited significant fibrosis when implanted in the peritoneal cavity of rats.63 Despite concerns about the immunogenicity, the advantageous roles of polycations prompted clinical trials for the treatment of diabetes with islet cells. The first successful clinical trial showed that the islet cells remained viable in the intraperitoneal cavity of a patient for 9 months.42 However, subsequent larger clinical trials have yet to show as promising results and have not led to commercialized products. As suggested in the previous section, the further coating of polycation layers with biocompatible polymers, such as alginate and PEG, may further improve the maintenance of cell viability over an extended time period.

The bulk pore size of a hydrogel is controlled by modifying the chemical structure of polymers and the cross-linking density of the gel matrix. The permeability of PEG hydrogels is controlled with the molecular weight of PEG molecules, because the difference of MWs alters the distance between cross-linking points and subsequently the swelling behavior and porosity of the gels. The PEG gels formed from cross-linking of polyethylene glycol diacrylate (PEGDA) with a molecular weight smaller than 8000 g/mol presented pores that effectively blocked the diffusion of immunoglobulin, albumin, ovalbumin, myoglobin, but not vitamin B12 (MW of 1.3 kDa). However, by increasing the molecular weight of the PEGDA to 20,000 g/mol, the gel became permeable to myoglobin (MW of 16.7 kDa), which is the size of some therapeutic proteins, but the gel remained impermeable to larger proteins, such as immunoglobulin.43 Islet cells encapsulated in these PEG gels and implanted in the peritoneal cavity of mice remained viable for more than 3 months and in rats for 1 month, without significant immune response.35 As expected, when PEGDA with MWs of 4000, 8000, or 10,000 g/mol were used to encapsulate β-cells, they did not inhibit the diffusion of small nutrients into the gel and allowed excellent cell viability. The encapsulated β-cell viability was dependent on maintaining cell–cell adhesion. The therapeutic efficiency of the encapsulated and aggregated β-cells remained high during transplantation into mice.25 These techniques to control the pore structure of hydrogel bulk may also be applicable to other materials, such as poly(vinyl alcohol) and poly(HEMA) gels.

The transport properties in hydrogels may be further enhanced with the parallel control of porous microstructure of the hydrogel and the inherent physiological environment of the cells. Strategies to prevent the influx of immune system proteins and allow the efflux of therapeutic molecules and wastes secreted from cells solely through the control of hydrogel pore size may not be sufficient for the long-term cell transplantation. Specifically, the molecular weight of nitric oxides secreted from macrophages is significantly smaller than therapeutic molecules such as insulin, endostatin, and nerve growth factor. Therefore, the pore size is large enough to allow the influx of the nitric oxides, which damages islet cells through the endoplasmic reticulum stress pathway.68, 69 These issues may be resolved with the use of scavenging cells and molecules. For example, erythrocytes encapsulated with islet cells protected the islet cells by scavenging the nitric oxides and preventing apoptosis.41 Separately, metallothionein and hemoglobin that effectively scavenge nitric oxides significantly increased the viability of transplanted islets.29, 31 Alternatively, islet cells genetically engineered to be resistant to cytokine-induced death were used to decrease the sensitivity of cells to the nitrous oxides.30 The preliminary results are promising, but there are still concerns about the use of genetically engineered cells in clinical trials.

CONTROL OF MECHANICAL PROPERTIES OF HYDROGELS

  1. Top of page
  2. Abstract
  3. INTRODUCTION
  4. CELL ENCAPSULATION PROCESS
  5. CONTROL OF HYDROGEL IMMUNOGENICITY
  6. CONTROL OF TRANSPORT PROPERTIES IN HYDROGELS
  7. CONTROL OF MECHANICAL PROPERTIES OF HYDROGELS
  8. CONCLUSION AND FUTURE DIRECTION
  9. References

The mechanical properties of cell-encapsulating hydrogels are highly important to ensure the persistent therapeutic efficacy of transplanted cells. Certain hydrogels may not present the desirable mechanical stiffness (resistance to deformation) and toughness (resistance to fracture) to structurally protect transplanted cells. Hydrogels may also lose their mechanical stiffness and structural integrity over time because of the several intrinsic and extrinsic factors exposing the encapsulated cells to the hostile immune system. These structural failures become critical problems as the hydrogels are implanted into tissues in which high mechanical stress is exerted, depicted in Figure 3. Various techniques are available to improve the mechanical properties and stabilities of hydrogels.

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Figure 3. The influence of external mechanical loading at implant site on the structural integrity of cell-encapsulating hydrogels may be mediated with mechanical properties of biomaterials. (a) Stiff biomaterials allow the maintenance of the hydrogel structure under compression. (b) Tough hydrogels play critical roles in preventing the hydrogel failure by slowing the crack propagation. [Color figure can be viewed in the online issue, which is available at www.interscience.wiley.com.]

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The mechanical properties of the gel are commonly controlled with the polymer concentration and molar ratio between polymers and cross-linking molecules. Increasing the polymer concentration and shortening the distance between the cross-links led to an increase of the mechanical stiffness.70 However, covalently cross-linked hydrogels become more brittle and susceptible to failure under mechanical loading exerted by neighboring tissue, as polymers are made increasingly stiff through increasing the cross-linking densities.71 In contrast, calcium cross-linked alginate hydrogels allow increases in both the stiffness and the toughness with increased cross-link density, unlike other hydrogels formed from covalent cross-linking. The transient nature of the ionic cross-links allow for this dual property control.44 It is likely that this accompanied increase of the gel toughness with stiffness allowed the successful use of these hydrogels during the transplantation of various drug-releasing cells. Mechanical properties of the hydrogels can also be modulated by reinforcing them with stiff inorganic particulates. Incorporating bioglass particles into the hydrogel significantly increased the gel stiffness. However, the capabilities of reinforced gels to support cellular viability in vivo remain to be tested.48 These approaches commonly alter the pore size of gel matrices as represented by the change of hydrogel swelling ratio. The changes of pore structure significantly influence biomolecular transports into and from hydrogel matrices, as discussed in “Control of transport properties in hydrogels” section. Therefore, it may be essential to develop a design principle to control hydrogel stiffness separately from the pore size of gel matrices for the independent control of hydrogel mechanics and transport properties.

The long-term mechanical stability of hydrogels is enhanced by modifying gel surfaces with polyelectrolyte complexes and cross-linkable molecules. These complexes formed with various polycations including poly-L-lysine, poly(ethyeleneimine), and PLO significantly increase the gel stiffness and toughness.45–47 Specifically, the use of PLOs to form polyelectrolyte complexes with alginate significantly enhances the durability of ionically cross-linked gels.40 This approach may also greatly enhance the durability of brittle covalently cross-linked gels. The formation of polyelectrolyte complexes also slows the ion-exchange process, which causes the gradual decrease of the mechanical stiffness of ionically cross-linked gels. Thus, calcium cross-linked alginate hydrogels reinforced by poly-L-lysine remain intact longer than plain alginate hydrogels in vitro. However, after implantation, their structural stability varies with the site of implantation, either subcutaneous or intraperitoneal, and the host organism, either canine or murine.46 The long-term mechanical stability of hydrogels may be further enhanced with semi-interpenetrating network gels and double-network models, which significantly improve the gel stiffness and toughness via secondary interactions between disparate polymers.72

The mechanical stiffness of hydrogels may furthermore influence the viability and function of encapsulated cells via integrin-ligand bonds. The stiffness of biomaterials regulates cellular activities of adherent cells, including proliferation, apoptosis, and differentiation. However, relatively little research has focused on the possibility that the mechanical properties of a 3D hydrogel matrix may regulate a cell's ability to secrete therapeutic molecules. The limited numbers of studies show debatable results. To an extent, encapsulating fibroblasts engineered to secrete VEGF into gels reinforced with bioglass enhanced the secretion level of VEGF. In contrast, encapsulating fibroblasts engineered to secrete bone morphogenetic proteins (BMP) into a softer PEG gel improved the secretion level of BMP.48, 49 A rigorous mechanistic study may elucidate the critical role of mechanical environments of cells in regulating secretion level of therapeutic drugs. This mechanistic study may eventually allow tuning of mechanical properties of cell-encapsulating hydrogels to an optimal level to improve the cell's ability to secreted drug molecules and also protect cells from external hostile environments.

CONCLUSION AND FUTURE DIRECTION

  1. Top of page
  2. Abstract
  3. INTRODUCTION
  4. CELL ENCAPSULATION PROCESS
  5. CONTROL OF HYDROGEL IMMUNOGENICITY
  6. CONTROL OF TRANSPORT PROPERTIES IN HYDROGELS
  7. CONTROL OF MECHANICAL PROPERTIES OF HYDROGELS
  8. CONCLUSION AND FUTURE DIRECTION
  9. References

Cell-encapsulating hydrogels are designed as a tool to improve the delivery of therapeutic drug molecules from cells. Cell viability and function to secrete therapeutic drug molecules is dependant on the cell encapsulation process, the level of immunogenicity, the transport control of biological molecules, and long-term structural stability of hydrogels. These hydrogel properties are improved by modification to the chemical structure and to the gelled microstructure. Certain approaches have significantly enhanced the therapeutic efficacy of cells in several preclinical trials, but further improvement is needed to attain the ultimate success in clinical trials.

The parallel control of the hydrogel structure, processing, and function may lead to further improvements in the therapeutic efficacy of cells. Specifically, chemical and physical properties of hydrogels often regulate the activities of adherent cells, including proliferation and differentiation, like a natural extracellular matrix.73, 74 In a similar manner, the hydrogel properties may also regulate the cellular secretion level of therapeutic drugs by activating the desired cell signaling and subsequently stimulating the gene expression level. Various nanoscale and microscale techniques will probably provide significant benefits in modulating individual properties of hydrogels to continuously control the cellular response.

Engineering the physiological environments of cell-hydrogel implants may also improve the clinical results of cell-based drug delivery. For example, the majority of cell-hydrogel constructs are implanted in the peritoneum cavity, but the undervascularized structure leads to hypoxic conditions, which limits the transport of oxygen, nutrients, and therapeutic molecules. Several tissue-engineering technologies to promote the capillary vascular network formation (e.g., delivery of angiogenic growth factors, transplantation of endothelial progenitor cells) may facilitate the transport of biological molecules from and into the hydrogels.

References

  1. Top of page
  2. Abstract
  3. INTRODUCTION
  4. CELL ENCAPSULATION PROCESS
  5. CONTROL OF HYDROGEL IMMUNOGENICITY
  6. CONTROL OF TRANSPORT PROPERTIES IN HYDROGELS
  7. CONTROL OF MECHANICAL PROPERTIES OF HYDROGELS
  8. CONCLUSION AND FUTURE DIRECTION
  9. References