Living Materials as Platform for New Technological Innovations

Current challenges impeding the successful realization of living materials are the demanding requirements for a viable environment for living organisms and biosafety concerns regarding the escape of genetically modified organisms into the environment (1). These issues led to design requirements, preferences and constraints that apply to our living material. These are general concepts, since the living material we propose can be used for many different applications. For each application, more specific criteria would apply. We go into more detail in our application scenario of a living material as a patch for burn & chronic wound healing (see below).

Requirements:
• Modularity: The living material must be capable of performing different functions by exchanging the type of incorporated genetically engineered E. coli.
• (Bio)safety: The living material must be designed in such a way that it is safe for both the users and the environment. This necessitates that bacterial leakage is prevented and the material is non-toxic and biodegradable.
• Viability: The living material must be a permissive environment for E. coli, allowing their survival and functioning in the material for a prolonged periods of time.
• Robustness: The living material must be able to withstand and function in a variety of conditions it might be subjected to in a real world application.

Preferences:
• Affordability: The production and maintenance of the living material is at a reasonable cost for the application.
• Ease of manufacturing & scalability: To allow widespread use of this new technological innovation, production should be straightforward and allow for upscaling of the manufacturing procedure to industrial scale.
• Sustainable: The living material and its production process must be environmentally-friendly.
• User friendliness: The material should be easy to use by the end user for a certain application.

Constraints:
• Considering the expertise of our department in E. coli biology compared to other bacterial strains, E. coli was the only logical candidate bacterial platform for our project.

Our material of choice for a living material is a hydrogel, a polymer network with water as dispersion medium. Hydrogels are inherently biocompatible, permeable and have a high water content. This makes them an ideal matrix for living materials (1). Their use as cell delivery vehicles (2) and scaffolds for tissue engineering (3) proves their suitability as a viable environment for micro-organisms. We quickly realized that our hydrogel has conflicting demands. It should prevent bacterial escape but also have an open porous structure to allow free diffusion of oxygen, nutrients, signaling molecules and proteins. Simple physical entrapment of the bacteria would therefore not suffice. We envisioned that the solution to this conflict is to actively tether E. coli tightly to our material. In this way, we prevent the diffusion of E. coli out of our gel while allowing the free transport of solutes. Our hydrogel-anchor combination should thus act as a sort of prison, keeping the E. coli bacteria alive and functional, but restricting their movement. The concept of Gelcatraz was born.

Figure 1. Structure of the 1.5 MDa MpIBP adhesin structure.

Adhesin

The first step to realizing Gelcatraz started with the identification of a suitable bacterial adhesin to anchor the E. coli bacteria to the hydrogel. Recently, a large 1.5 MDa adhesin was found in the Antarctic bacterium Marinomonas primoryensis. This large adhesin, named M. primoryensis ice-binding protein (MpIBP), is 0.6 µm long and consists of more than 130 domains. Its structure is shown in Figure 1. 120 of these domains encode for the same immunoglobin-like β-sandwich that rigidifies upon binding of Ca²⁺ and acts as extender projecting the binding domains positioned on top of the chain into the medium (4).

In our design process, we envisioned that we could use the sugar binding domain of MpIBP, to bind to a sugar-based polymer. In this way, we can tether the E. coli bacteria to our hydrogel. On the other parts of the wiki, we will refer to the shortened version of MpIBP including the sugar binding domain as ‘adhesin’.

In its native host, the MpIBP adhesin is exported out of the cell using the single-step type I secretion machinery. The type I secretion of the adhesin requires four protein domains. Hemolysin A (HlyA) is a protein domain C-terminal to the adhesin protein, signaling for type I secretion. Hemolysin B (HlyB), Hemolysin D (HlyD) and TolC are the other proteins which are responsible for forming a trans-periplasmic export channel upon interaction with the HlyA domain, as shown in Figure 1 (5). In type I secretion, all proteins passing the β-barrel pore of TolC, a constitutively expressed outer membrane protein(6), must remain unfolded (4). The RIM and RIN domains of the adhesin (see Figure 1) are the only domains that fold independent of calcium. Calcium is naturally present in millimolar concentrations in sea water, the natural environment of M. primoryensis. The entire adhesin is translocated to the extracellular space where the domains fold, but the RIM and RIN domain are already folded in the cytosol and prevent translocation through the TolC protein. The RIM protein fits inside the TolC pore, with the RIN being a bulky anchor preventing total release of the adhesin. The RIM & RIN domain act as a TolC β-barrel plug (4). The translocation process of the adhesin is shown in figure 2. After translocation, the HlyB/D complex is recycled. In our design, we want to use this ingenious system to express the adhesin on the surface of E. coli. However, the native adhesin is too large to be produced in E. coli. To avoid problems with expressing the adhesin in E. coli, the amount of RII repeats in the gene encoding the adhesin was reduced to 15.

Figure 2. Translocation mechanism of MpIBP adhesin.

Hydrogel

To find an appropriate hydrogel for our adhesin to bind to, we used the fact that this sugar binding domain has its highest affinity for glucose (7). Our hydrogel must be biocompatible, it must be non-toxic to E. coli. The microstructure of the hydrogel must be open and porous, allowing efficient and fast nutrient, oxygen and solute exchange. The hydrogel must be mechanically stable to prevent bacterial leakage. The hydrogel must not degrade over a sustained period of time and synthesis should be relatively straightforward and reproducible.

Dextran is a polymer that consists of α-1,6-linked D-glucopyranose (glucose) residues. This glucose-based nature makes dextran an ideal polymer for our MpIBP adhesin to bind to. Dextran is non-toxic, has been investigated for tissue engineering applications, and is already used in the clinic as blood plasma substitutes (8). A hydrogel from dextran polymer chains can be synthesized using radical polymerisation of glycidyl methacrylated dextran (9). By adapting the protocol of Plieva et al.(9), we designed a supermacroporous dextran hydrogel with pores ranging from 1 to 100 micrometer that meets our requirements. This unique morphology allows for the rapid and non-restricted mass-transport of solutes of virtually any size (9). Figure 3 illustrates a schematic representation of the living material we envision.

Figure 3. Schematic representation of the envisioned biomaterial.

Testing Our Living Material Platform

To demonstrate the performance and behaviour of our designed living material platform, we designed and carried out multiple experiments. A detailed analysis of the results can be seen here. We started with reproducing the hydrogel synthesis conditions described in the paper by Plieva et al.(9). When this yielded mechanically unstable hydrogels, we examined different combinations of degree of substitutions of dextran, equivalents of the radical initiator system, and preparation temperature to identify suitable conditions that yield reproducible hydrogels. See here for more information.

Subsequently, Scanning Electron Microscopy experiments were performed to find the optimal hydrogel polymerisation conditions that result in the desired macroporous structure. After attempting different hydrogel synthesis conditions, the required conditions were identified.

To demonstrate the functionality of the dextran-adhesin combination, we first investigated the dextran-binding capabilities of the isolated MpIBP carbohydrate-binding domain via an assay based on dextran binding affinity to Sephadex beads, which can be used to pull dextran-binding proteins from solution. A schematic overview and global explanation of this experiment can be seen in figure 4. More details about this experiment and the results can be seen at the registry page of our carbohydrate-binding domain biobrick, BBa_K2812000.

Figure 4. Assay to examine the dextran binding capabilities of the carbohydrate-binding domain. The beads present in step 1 are dextra-based Sephadex beads to which dextran-binding proteins can bind and subsequently these beads can be used to pull dextran-binding proteins from solution.

Next, a washing experiment was designed to investigate the binding of adhesin expressing E. coli BL21 (DE3) bacteria to 380 µL columnar-shaped hydrogels under different circumstances. A schematic overview and global explanation of this experiment can be seen in figure 5. Click for more detailed experimental conditions and results. During these experiments we switched in our design from E. coli BL21 (DE3) to E. coli BLR, a strain known known for more stable expression of proteins with many repeats, in an attempt to obtain more consistent results.

Figure 5. Schematic illustration of the washing experiment set-up to investigate the effect of the adhesin in tethering the bacteria to the hydrogel. 1) Culture 2) Induction, seeding in gel 3) Washing 4) Crushing 5) Samples taken.

The washing and crushing experiment was time consuming and error sensitive. Therefore, a new experiment was devised to investigate the effect of the adhesin in retaining the E. coli bacteria inside the hydrogel. The central component in this experiment is a cubic hydrogel containing a cavity, which we fabricated using polydimethysiloxane fabrication. The chamber allows bacteria to be loaded inside the hydrogel and subsequently, a leakage experiment was performed to quantify bacterial leakage over time. A schematic illustration can be seen in figure 6. Click for more detailed experimental conditions and results With this new prototype hydrogel, more consistent and robust results were obtained.

Figure 6. Schematic illustration of the leakage experiment set-up to investigate the effect of the adhesin in tethering the bacteria to the hydrogel. 1) Pre-culture with induction 2) Inject culture into cavity 3) Gels soaked in culture medium 4) Gel transferred into petri dish 5) Bacteria injected into cavity 6) Medium added. Afterwards, samples are taken periodically.

Fighting against Infections of Burns and Chronic Wounds

To illustrate the potential of living materials, we decided to develop it into a living skin patch capable of the continuous secretion of bacteriocins to combat bacteria responsible for infections in patients suffering from burns and chronic wounds. For a detailed analysis of the background of the problem and our interviews with experts and stakeholders please see our Human Practices section and Applied Design section.

As mentioned above, this specific application scenario leads to additional requirements, preferences and constraints.

Requirements:
• Safety & Biocompatibility: The bacteriocin secreting patch should be safe to use for patients and healthcare professionals. It should be easy to sterilize prior to seeding with bacteria to avoid the introduction of other bacteria near the wound.
• Effectivity: The effect of the patch should be limited to infection causing bacteria, with Staphylococcus aureus as our primary target
• User friendliness: The patch must be easy to use for healthcare professionals such as doctors and nurses. It must also be comfortable for patients.
• Selectivity: The bacteriocins secreted by the patch must only be active against the specific strains of bacteria causing wound infections.

Preferences:
• Modularity: Ideally our patch can target different bacterial strains by simply changing the bacteriocin encoding gene.
• Longevity: Our patch should be able to perform its function for approximately 3 to 4 days, reducing the number of wound dressing changes patients experience.
• Storability: The patch should have a long shelf-life and should be straightforward to store to compete with already existing treatments and to avoid availability issues.
• Approval: The base material of the patch (the hydrogel) should ideally be already approved by the U.S. Food and Drug Administration (FDA) and European Medicines Agency such that only the polymerisation and integration of living organisms requires investigation and approval by regulatory institutions to accelerate implementation.

Constraints:
• The bacteriocins secreted by the patch must not be toxic to E. coli

Our dextran hydrogel is very suitable for wound healing. Our hydrogel is stable at physiological conditions for months and can be stored in a dried state and re-swollen prior to use, avoiding availability problems (9). Dextran is already approved by the FDA for clinical use for plasma volume expansion, antithrombotic agent and peripheral flow promotion (10). Furthermore, the hydrogel can be easily sterilized using radiation (11). For this application scenario, FDA approval for the polymerized methacrylated dextran, possible degradation products, and the integration of engineered E. coli bacteria is required.

To target Staphylococcus aureus, a Gram positive bacteria, we use lysostaphin. Lysostaphin is a zinc endopeptidase that cleaves the pentaglycine cross-bridges present in the bacterial cell wall of S. aureus. Lysostaphin exhibits its lytic effect not only against regular S. aureus, but also against methicillin-resistant S. aureus (MRSA), vancomycin-resistant S. aureus and S. epidermidis. Recently, it was demonstrated that hydrogels delivering lysostaphin outperform prophylactic antibiotic treatment and soluble lysostaphin treatment in murine femur fracture models (12). The iGEM teams HIT-Harbin 2012 and Stockholm 2016 have used lysostaphin before in their projects to target S. aureus. However, both teams relied on cell lysis to free the lysostaphin. For our application, this approach is not suitable as releasing bacterial cell lysate into the wound will only act as food for the bacteria causing the infection. Moreover, cell lysis does not offer a solution for sustained release of lysostaphin. The solution to this problem is to make lysostaphin secretable. This will allow sustained release of the active enzyme and ensures that the lysostaphin reaching the wound is of high purity.

Enabling secretion

Different strategies for secretion of recombinantly expressed proteins exist (13). To minimize the burden on our modified E. coli bacteria that are already expressing MpIBP adhesin, secretion of lysostaphin should use the same type I secretion system. To achieve this, the C-terminal HlyA signal peptide must be fused to lysostaphin. In designing the construct encoding for such a fusion protein, we first codon optimized the HIT-Harbin 2012 truncated lysostaphin construct and fused it to HlyA via a thrombin linker. Secretion of lysostaphin circumvents the use of cell lysis, improving the biokiller device developed by HIT-Harbin 2012. The thrombin linker was chosen with other iGEM projects in mind, as it can be readily cleaved by the thrombin enzyme. For protein detection and purification purposes, we added a His-tag C-terminal to the HlyA domain.

After assembly of this construct, protein expression and functional experiments in collaboration with the PAMM foundation, the regional center for infectious diseases and pathology of South East Brabant in the Netherlands, were performed. To prove the functionality and secretion of the lysostaphin in collaboration with PAMM, we devised a number of experiments. First, cell lysate was tested on S. aureus to show the functionality of lysostaphin. This experiment was followed by testing the of the medium, to see whether lysostaphin is being secreted by the cells. Finally, as ultimate test of the secretion and functionality of the lysostaphin, co-cultures were grown with S. aureus and our induced bacteria. All these experiments were performed with Streptococcus agalactiae as a negative control, a Gram positive bacteria similar to S. aureus, but not susceptible to lysostaphin. The HlyA domain itself is also used as a control, to eliminate HlyA as source of lytic activity against S. aureus. Click for more detailed experimental conditions and results. Also, see our part BBa_K2812004 for more information, including the improvement upon BBa_K748002.

Figure 7. A schematic of our wound healing patch.

During our conversations with experts on the topic of burn and chronic wounds, we learned that besides S. aureus, Pseudomonas aeruginosa is also a major cause of problems in wound care of patients with burn and chronic wounds. As our lysostaphin-HlyA has a modular design, we changed the coding sequence of lysostaphin to that of pyocin S5, generating a new biobrick in addition to lysostaphin. Pyocin S5 is a bacteriocin against P. aeruginosa(14), enabling our patch to target both S. aureus and P. aeruginosa simultaneously. This demonstrates the modularity of our patch, of which the concept is illustrated in figure 7. We also learned about the importance of a kill switch and included this in our design. Please see our Applied Design section for the kill switch and for an analysis of the potential impact of our bacteriocin secreting patch.

Model results were used to adapt the hydrogel shape to suit our wound healing application. We modelled which patch thickness would be feasible to create a patch that is as thin as possible, but also allows enough lysostaphin to reach the skin to kill S. aureus using experimental input as lysostaphin production rate and hydrogel paramters such as microstructure. Our model showed that a patch with a thickness of the order of magnitude of approximately a millimeter allows a high enough concentration of lysostaphin to kill S. aureus. Subsequently, we adapted our hydrogel patch design to synthesize a hydrogel patch with a thickness of 0.75 mm. We could show that this hydrogel could be easily created and yielded flexible, stable hydrogels. For our wound healing application, we could develop a relevant prototype due to the modeling results.

Summary of our design

The core of our project consists out of three DNA constructs; the secretable lysostaphin, the type I secretion system proteins HlyB/D and the MpIBP adhesin. To comply with the engineering principle of modularity, each of these three constructs is under to control of a different promotor. This allows individual manipulation of each component. Additionally, to maintain the plasmids during division of the bacteria, each construct has an unique origin of replication. To confirm uptake of all three plasmids, each construct has a different antibiotic resistance. The three main construct designs of our project can be seen in figure 8. We used experimental data to model the ideal prototype hydrogel for the wound healing application, and subsequently used these outcomes to optimize our patch design and synthesis.

Figure 8. The three core constructs of the design of our project.

We achieved modularity by using unique promotors and vectors for each component allowing individual manipulation and optimisation the integrated system. We used prototyping to generate different shapes and sizes of our hydrogel and used PDMS fabrication to improve our design of the hydrogel. By performing all experiments at least in duplo and repeating them multiple times we ensured the reproducibility of our results. Via integration of our conversations we had with experts for Human Practices and Applied Design parts of our project, we implemented safety as a key feature in our design.

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