Engineering a clinically-useful matrix for cell therapy

Abstract

The design criteria for matrices for encapsulation of cells for cell therapy include chemical, biological, engineering, marketing, regulatory, and financial constraints. What is required is a biocompatible material for culture of cells in three-dimensions (3-D) that offers ease of use, experimental flexibility to alter composition and compliance, and a composition that would permit a seamless transition from in vitro to in vivo use. The challenge is to replicate the complexity of the native extracellular matrix (ECM) environment with the minimum number of components necessary to allow cells to rebuild a given tissue. Our approach is to deconstruct the ECM to a few modular components that can be reassembled into biomimetic materials that meet these criteria. These semi-synthetic ECMs (sECMs) employ thiol-modified derivatives of hyaluronic acid (HA) that can form covalently crosslinked, biodegradable hydrogels. These sECMs are “living” biopolymers, meaning that they can be crosslinked in the presence of cells or tissues to enable cell therapy and tissue engineering. Moreover, the sECMs allow inclusion of the appropriate biological cues needed to simulate the complexity of the ECM of a given tissue. Taken together, the sECM technology offers a manufacturable, highly reproducible, flexible, FDA-approvable, and affordable vehicle for cell expansion and differentiation in 3-D.

Introduction

An ever-increasing number of hydrogels and fibrous scaffolds based on natural and synthetic polymers are being developed for the expansion of primary cells and progenitor cells in three dimensions (3-D).1–4 Among the naturally-derived biomaterials is an in situ-crosslinkable synthetic mimic of the extracellular matrix (sECM) that is based on chemically-modified hyaluronan (HA).5 HA is a ubiquitous non-sulfated glycosaminoglycan with critical roles in morphogenesis,6 and chemically-modified HA derivatives have generated considerable interest as biomaterials for tissue engineering.7,8 This versatile sECM offers user-controllable composition and compliance, to which a variety of growth factors or matricellular proteins may be added.9,10 The elastic moduli of hydrogels can be tailored to match the cell type to achieve optimal regeneration of a new tissue.11,12

Design criteria for ECM mimetics.

For a biomaterial to substitute for the ECM, it needs to recapitulate the principal functions of the natural ECM in orchestrating cell proliferation, migration, differentiation, angiogenesis and invasion.13,14 Recreating the interwoven set of biochemical and mechanical cues in the cellular microenvironment2,15 and allowing full experimental versatility in fabrication methods for engineered microenvironments1,16,17 requires consideration of a variety of design criteria. As summarized in Figure 1, the performance criteria that should be met by a biomimetic ECM substitute should include: (1) experimentally variable composition, (2) experimentally variable compliance, (3) controllable biodegradability in vitro and in vivo, (4) multiple physical forms for in situ crosslinking during cell encapsulation and fabrication, (5) batch to batch consistency, (6) ease of use at physiological temperature and pH, (7) transparency for ease of visualization, (8) compatibility with high throughput screening (HTS) platforms, and (9) translational potential. It is important that the sECMs are modular in nature, allowing the user to add soluble factors, attachment peptides, matricellular proteins, crosslinkers, and macromomers in whatever combinations9 are dictated by the need to alter experimental parameters for testing hypotheses for a given cell type.

Most recently, we have underlined the importance of including the end-user needs as design criteria for materials to be used for cell therapy5 and for development of drug evaluation tools.18 These additional design criteria (Fig. 2) are derived from the truism that “nothing from the lab ever reaches a patient unless it is commercialized.” Working backwards from this irrevocable reality, we recognized that commercialization—at least in the United States—requires significant financial commitment to developing the technology and to recruiting a team capable of reducing the technology to practice. Investor interest is highest for products for which a large market can be readily penetrated. Specifically, for biomaterials to be used in regenerative medicine, several key parameters must be been satisfied.

Deconstructing the ECM: A Chemist's Solution to a Bioengineering Problem

The native ECM is a heterogeneous network of soluble and fibrous proteins and glycosaminoglycans (GAGs), which is maintained by both covalent and non-covalent interactions (Fig. 3). In analogy to concrete, the ECM contains fibrillar collagen (rebar), proteoglycans (pebbles), hyaluronan (cement), and other minor components. To engineers and chemists, recapitulating the composition and structure of the ECM is a daunting yet irresistible challenge. To this end, we deconstructed the ECM into modular units containing one, two, or several chemically modified macromonomers that could be covalently crosslinked into a hydrogel network (Fig. 4).

The one-component solution.

In the last two decades, a host of new chemistries have been developed, resulting in a series of chemically-modified HA medical products that have addressed clinical needs with different levels of success. 8,19 In addition, chemical modification techniques have allowed the modification of the surfaces of implantable devices and biomaterials with HA.20

Chemically-modified HA derivatives were recently grouped into two categories.21 When the modification resulted in a final form that could not form new chemical bonds in the presence of cells or tissues, and could only be manipulated by fabrication into different physical forms, these were classed as “monolithic” derivatives. In contrast, when new covalent bonds could be formed in the presence of cells or tissues, enabling a change in physical form during in vivo or in vitro biological use, these were classed as “living” derivatives. Thus, an in situ-crosslinkable material for adhesion prevention or cell encapsulation would be a “living” modification. This review focuses on living HA technologies and products and their uses for 3-D cell culture, wound repair, and tissue engineering.22–24

A synthetic extracellular matrix (sECM).

A novel approach to the preparation of a covalently-crosslinked sECM was recently described.24,25 The sECM material may be crosslinked in situ in the presence of cells in order to provide an injectable cell-seeded hydrogel for tissue repair26 or without cells for the controlled, localized delivery of drugs or growth factors.27–29 Thiol-modified glycosaminoglycans and proteins are living macromonomers that can be crosslinked with biocompatible polyvalent electrophiles to form biocompatible, bioerodable hydrogels and porous sponges. This new technology offers a powerful set of new tools for culturing primary cells in three dimensions (3-D). Indeed, the applications of this technology include in vitro and in vivo growth of 3-D cellularized, vascularized tissues using films, sponges, and hydrogels.

The modular, clinically versatile and readily-manufactured sECM was made possible using hydrazide reagents that contain disulfide bonds,30 affording thiol-modified proteins and GAGs can be prepared and crosslinked into hydrogels under cytocompatible physiological conditions.31 The thiolated macromonomers may be subsequently crosslinked in air only to a hydrogel, which may be dried as a thin film or lyophilized to produce a porous sponge. Alternatively, crosslinking with difunctional electrophiles can be accomplished, in the presence or absence of cells, to give injectable and biocompatible hydrogels.23,26 Different physical properties and rates of biodegradation can be obtained by controlling several parameters for the biocompatible HA hydrogels32: (1) molecular weight of starting HA employed; (2) percentage DTPH modification; (3) concentration of HA-DTPH; (4) molecular weight of PEGDA; and (5) ratio of thiols to acrylates.

Photo-crosslinked HA: An alternative living polymer building block.

Methacrylate and methacryamide derivatized HA are living macromonomers that can be seeded with cells and photocrosslinked into hydrogels with either UV or visible light plus an initiator.33–36 The potential uses for these in situ photopolymerized hydrogels include prevention of adhesions and tissue regeneration; most recently, photopolymerized HA methacrylate was used for the expansion and differentiation of human embryonic stem cells.37 For example, chondrocytes encapsulated in photocrosslinkable HA methacrylate retained the chondrocytic phenotype and synthesized cartilage matrix in vitro, and accelerated healing in an osteochondral defect model in vivo.38 The poor cell attachment to these in situ photopolymerized HA hydrogels can be significantly enhanced by RGD modification.33

Using the sECM platform in vivo.

The thiol modified HA component, first HA-DTPH but subsequently a carboxymethylated HA known as CMHA-S, was readily cross-linkable and yielded biocompatible hydrogels that did not support cell attachment and proliferation. Crosslinking was accomplished slowly via disulfide bond formation in air, or with better control of crosslinking time and hydrogel stiffness using cytocompatible bivalent cross-linkers such as polyethylene glycol diacrylate (PEGDA),39 PEG divinyl sulfone40 or PEG bis(haloesters) and bis(haloamides).41 This one-component gel has proven utility in scar-free wound healing, post-surgical adhesion prevention, and stem cell expansion. Specifically, the one-component gels were used in vivo for vocal fold repair,42 to reduce post-operative intra-abdominal and pericardial adhesions,43,44 to eliminate postsurgical tendon adhesions,45 to reduce airway stenosis,46 to improve healing following endoscopic sinus surgery 47 and to promote healing following middle ear injury.48

An example of the use of the PEGDA-crosslinked CMHA-S is illustrated in Figure 5. Using a murine full-thickness wound repair model, we evaluate a one-component sECM composed of only the crosslinked carboxymethylated HA derivative. Bilateral injuries were formed on the flanks of each animal, allowing each animal to act as its own control. One site was left untreated, while a crosslinked hydrogel film of PEGDA-crosslinked CMHA-S was applied to the contralateral wound. A sheet of Tegaderm was laid across both wounds to prevent dessication and reduce microbial invasion, and a bandage was applied to retain the dressings in place. After six days, wound area was quantified, and then animals were sacrificed to perform histology. The treated wounds were over 45% re-epithelialized after six days, while the untreated wounds were only 12% healed. Moreover, wound filling, epidermal thickness, dermal thickness, and vascularization were also significantly enhanced in treated animals.

The two-component solution.

To obtain materials on which cells could readily attach and spread, we incorporated thiol-modified gelatin (Gtn-DTPH),26,49 RGD peptides50 and cysteine-containing domains of recombinant human fibronectin51 into the hydrogel formulations. This produced the two-component gels shown in the middle panel of Figure 2. Alternatively, simple addition of collagen provided a non-contracting 3-D hydrogel suitable for culture of human fibroblasts.52 Such two-component hydrogels were tested with excellent results in three-dimensional (3-D) cell culture and tissue engineering applications.23 Two quite different physical forms of the sECM were evaluated in vivo, as illustrated below.

First, we evaluated the use of an injectable, cell-free, two-component sECM hydrogel for repair of an osteochondral defect in a rabbit patellar groove defect (Fig. 6).53 In untreated control defects, extensive scar tissue accumulated after twelve weeks. However, simply filling the defect with the sECM hydrogel resulted in regeneration of both new osteal and chondral tissues. Apparently, host-derived cells invaded the compliant hydrogel, remodeled it with matrix metalloproteinases, and secreted a new native ECM specific to either the bone or cartilage environment. Clearly, the simple HA-gelatin hydrogel is neither providing structurally or biochemically complete “instructions” for bone or for cartilage regeneration. Nonetheless, a biodegradable, osteo- and chondroconductive environment was provided by the hydrogel. Indeed, a more robust remodeling was accomplished by the addition of autologous MSCs to the hydrogel, suggesting that localized cell delivery and retention can add an important dimension to the regenerative process.45 Still, the fundamental role of the hydrogel is to enhance the natural biological repair processes mediated by endogenous cells.

Second, we tested the hypothesis that the stiffness of the ECM substitute was important in the reparative context. For this study, we employed a rat femoral defect model, using either a hydrogel form of the cell-free two-component sECM Extracel™, or a lyophilized sponge of Extracel™.54 We also tested whether the gels and sponges containing human demineralized bone matrix (DBM), a xenogeneic bone-derived growth factor mixture, would further enhance bone repair. As shown in Figure 7, the untreated control injury showed little new bone formation after eight weeks. In contrast, we found that the stiffer sECM sponge was superior at all time points to the softer sECM hydrogel in generating new bone. In addition, the addition of DBM to either material gave a statistically significant improvement in healing.54 Perhaps unexpectedly, stiffness was more important than growth factors; a sponge without DBM was more effective than a gel with DBM. In the end, the optimal treatment consisted of the Extracel” sponge plus DBM, which showed >95% closure of the defect and filling of the defect with healthy bone tissue after eight weeks.

Adding complexity one step at a time-sulfated GAG-containing sECMs.

Engineered tissues ultimately require a blood supply, so that thiolated heparin (HP-DTPH) and chondroitin sulfate (CS-DTPH) macromonomers were developed to provide the biomimetic equivalent of heparan sulfate proteoglycans. We demonstrated that these hydrogels provided spatiotemporal control of growth factor release over a time course of weeks to months. Basic fibroblast growth factor (bFGF) was released from an HA-based sECM in vitro with a half-life of one month;27 when released from a CS-based sECM in vivo, bFGF showed dose-dependent acceleration of re-epithelialization and improved formation of vascularized dermal tissue in the repair of large acute wounds in diabetic mice.55 Dual delivery of vascular endothelial growth factor (VEGF) with bFGF,56 with angiopoetin-1 (Ang-1), and with keratinocyte growth factor (KGF)57 led to enhanced neovascularization in vivo. For example, Figure 8 shows mature vasculogenesis in a mouse ear pinnae model when both KGF and VEGF from a heparan proteoglycan-mimetic sECM hydrogel, similar to what is seen in native vasculature. In contrast, VEGF-only hydrogels generated only immature leaky angiogenic sprouts.

The Translational Challenge

The translational challenge in replicating the native ECM environment requires co-optimization of biological, structural, and practical design parameters. Too often, biologists and engineers have minimized the importance of the financial, production, and regulatory constraints when developing a biomaterial for reparative or regenerative use in the clinic. This has resulted in many false starts. In contrast, a marketing approach driven by the needs of the end-users places these crucial and limiting pragmatic factors as part of the initial design criteria. In this way, we have advocated using the minimum number of components necessary to allow cells to rebuild a given tissue.

The message of this rearrangement of the discovery process for new biomaterials is summarized in Figure 9. The physiological and structural aspects of rebuilding tissues from cells—from whatever source—are immensely complicated. We must embrace this complexity, but it cannot dominate the design of a clinical material. Moreover, one formulation cannot fulfill every need, from stem cell expansion and delivery for neural tissues to creating a structurally robust material for large-scale osteochondral defect repair. We should engineer flexibility, so that a single suite of approvable materials can be exploited for multiple physician uses in a variety of clinical niches. Finally, for regenerative medicine to succeed as a reality, we must deliver simplicity. The penetration of new technology into any market is not determined only by improvement in patient outcome —the true limiting factors are cost, familiarity, and ease of use. We believe that the living sECM technology addresses a balance between the physiological and the practical requirement, and that its uses for clinical, veterinary, and laboratory research needs will continue to expand in the decades to come.

Abbreviations

abbreviations

=

Figures and Tables

Figure 1 Performance criteria for sECMs for 3-D cell culture and translational utility.

Figure 2 Additional design criteria for a synthetic extracellular matrix optimized for reparative medicine.

Figure 3 The native extracellular matrix is a heterogeneous networked complex of proteins and glycosaminoglycans that is maintained by multiple covalent and non-covalent interactions.

Figure 4 Deconstruction of the ECM to one, two, and multi-component modular, living sECMs.

Figure 5 (A) Wound repair model and evaluation of a one-component sECM composed of only a crosslinked HA derivative. (B) Improved histological outcomes include wound filling, production of a vascularized dermis, and regeneration of a new epithelium complete with stratum corneum.

Figure 6 Cartilage repair in an osteochondral defect with a cell-free two-component sECM hydrogel. Left, rabbit patellar groove defect with no treatment results in extensive scar tissue after three months. Right, filling the defect with Extracel™ hydrogel results in regeneration of both new osteal and chondral tissues by invasion of the sECM by host-derived cells.

Figure 7 Bone repair in a femoral defect model with a cell-free two-component Extracel™ lyophilized sponge containing human demineralized bone matrix (DBM). Above, untreated control injury results in little defect closure (autoradiogram) and extensive scar tissue (histology). Below, optimal treatment with Extracel™ sponge plus DBM shows complete closure of the defect and filling of the defect with healthy bone tissue.

Figure 8 Mature vasculogenesis in mouse ear pinnae occurs with release of both KGF plus VEGF from Extracel™-HPG hydrogels (right) as compared to native vasculature (left), while VEGF-only hydrogels (middle) generate only immature leaky angiogenic sprouts.

Figure 9 The translational challenge

Acknowledgements

We thank the Utah Centers of Excellence Program for the primary financial support for this work. Additional resources from NIH Grant R01 DC4336 from the National Institute of Deafness and other Communicative Disorders, and from the NSF Frontiers in Innovative Biological Research Grant EF0526854 contributed to the conceptualization and implementation of the sECMs described herein.