Biomimetic peptide conjugates as emerging strategies for controlled release from protein-based materials

Abstract

Biopolymers, such as collagens, elastin, silk fibroin, spider silk, fibrin, keratin, and resilin have gained significant interest for their potential biomedical applications due to their biocompatibility, biodegradability, and mechanical properties. This review focuses on the design and integration of biomimetic peptides into these biopolymer platforms to control the release of bioactive molecules, thereby enhancing their functionality for drug delivery, tissue engineering, and regenerative medicine. Elastin-like polypeptides (ELPs) and silk fibroin repeats, for example, demonstrate how engineered peptides can mimic natural protein domains to modulate material properties and drug release profiles. Recombinant spider silk proteins, fibrin-binding peptides, collagen-mimetic peptides, and keratin-derived structures similarly illustrate the ability to engineer precise interactions and to design controlled release systems. Additionally, the use of resilin-like peptides showcases the potential for creating highly elastic and resilient biomaterials. This review highlights current achievements and future perspectives in the field, emphasizing the potential of biomimetic peptides to transform biopolymer-based biomedical applications.

Graphical Abstract

1. Introduction

Controlled release systems are essential components in modern biomedical applications, particularly in drug delivery. These systems are designed to deliver therapeutic agents in a regulated manner over a specified period, thereby enhancing the efficacy of treatment and minimizing side effects (Li et al., 2022; Liu et al., 2024). By controlling the rate, duration, and spatial release of drugs, these systems can maintain optimal drug concentrations at the target sites, reduce dosing frequency, and improve patient compliance. The importance of controlled release systems spans various medical fields, including cancer therapy, chronic disease management, and regenerative medicine.

Protein-based materials are promising in these applications due to their natural compatibility with biological systems, their degradability, and their ability to support cell growth and tissue integration. When used as carriers, protein-based materials can provide a favorable environment for therapeutic agents, enhancing the stability and bioactivity of encapsulated molecules (Luan et al., 2024). Additionally, the structural versatility of proteins, such as collagen, silk fibroin, and fibrin allows for material customization to enhance cell interactions and facilitate sustained therapeutic effects (Zhu et al., 2023). However, therapeutic agents loaded within these materials often face challenges with rapid, uncontrolled release, which can lead to a short-lived therapeutic effect. This rapid release, commonly referred to as ‘burst release,’ frequently limits the efficacy of bioactive agents, necessitating repeated administrations, which can increase treatment costs and patient inconvenience. Moreover, protein-based materials exhibit batch-to-batch variability, potential immune responses, and inconsistencies in mechanical properties, complicating their application in controlled release systems. To address these issues, the incorporation of biomimetic peptides can enhance the controlled, sustained release of bioactive agents. These peptides interact with native proteins in a manner that can help stabilize release kinetics, thereby maximizing therapeutic benefits and reducing the downsides of uncontrolled release (Hu et al., 2015; Chattopadhyay et al., 2016; Wang et al., 2019; Liang et al., 2021; Chattopadhyay et al., 2022; Promsuk et al., 2024).

The consequences of uncontrolled release of bioactive proteins from hydrogels can be significant. Platelet-rich plasma (PRP) contains multiple growth factors that are released by 95% within an hour after injection. The growth factors then diffuse into the tissue fluid and are quickly cleared into the circulation, failing to sustain their effects on cell proliferation and angiogenesis (Liu et al., 2017; Chang et al., 2018). Another example is the development of chimeric antigen receptor (CAR) T-cell-encapsulated materials for solid tumor treatment. In this case, proinflammatory cytokines, such as IL-2, TGF-β, or IL-15 are incorporated into hydrogels or particles to create a stimulatory microenvironment that activates the CAR T-cells to infiltrate and act against the solid tumor (Liu et al., 2018; Zhou et al., 2022). If these cytokines are rapidly depleted due to uncontrolled burst release, the CAR T-cells have a short active period, leading to ineffective treatment. Furthermore, uncontrolled burst release of these cytokines may lead to a cytokine storm, causing severe systemic inflammation and potential harm to the patient. Therefore, strategies to control the release rate of bioactive proteins are essential.

The growing field of biomimetic peptides has opened up new possibilities for controlled release systems in protein-based materials. Biomimetic peptides, designed to mimic the structure and function of natural peptides, interact with the corresponding native proteins. These interactions are crucial for sustaining the release of bioactive molecules, ensuring their delivery in a controlled manner over time. In various protein-based materials, such as collagen, elastin, silk, and fibrin, biomimetic peptides have demonstrated the ability to form stable complexes. These complexes can modulate the release rates of bioactive molecules, making them highly effective in drug delivery and tissue engineering applications. For instance, in collagen-based systems, peptides that mimic domains of natural collagen can enhance the stability and controlled release of growth factors (Chattopadhyay et al., 2022). Similarly, silk-fibroin mimetic peptides, which replicate the repeating sequences of natural silk fibroin, can be engineered to release a bio-functional protein at a tunable rate from a hydrogel (Promsuk et al., 2024).

The interactions between biomimetic peptides and native proteins often involve non-covalent interactions, such as hydrogen bonding, hydrophobic interactions, and ionic bonding. These interactions help maintain the structural integrity of protein-based materials while enabling the sustained and controlled release of bioactive molecules. By understanding and harnessing these interactions, researchers can design more efficient delivery systems to improve therapeutic outcomes.

In addition to the benefits of controlled release, covalently conjugating biomimetic peptides to bioactive proteins offers advantages for manufacturing. Genetic engineering techniques can produce both peptides and proteins in one go, simplifying the production process. This approach not only reduces manufacturing complexity but also facilitates easier scaling for industrial production. Exploiting genetic engineering makes it feasible to produce large quantities of these conjugates consistently, making the technology more accessible for commercial applications (Lin et al., 2024).

This review aims to provide a comprehensive overview of strategies for achieving controlled release from protein-based materials through biomimetic peptides, focusing on the unique contributions of these peptides to stabilizing and tuning release profiles. Unlike other existing reviews, which center on the motif design of protein-mimetic peptides for diverse applications (Zhang et al., 2018) or emphasize self-assembled peptides in imaging and cancer biotherapy (Luan et al., 2023), this review highlights the specific potential of biomimetic peptides to control bioactive molecule release by engaging in sustained interactions within protein-based delivery systems.

This review addresses various types of protein-based materials, including collagen, elastin, silkworm silk fibroin, spider silk, fibrin, keratin, and resilin (Figure 1), examining their inherent properties, roles in controlled release systems, and the ways in which biomimetic peptides enhance their performance. Additionally, genetic engineering enables the production of biomimetic peptides attached to protein cargos, simplifying manufacturing and allowing for scalable production in industrial applications. This comprehensive approach highlights emerging trends, challenges, and future directions, providing insights into the promising combination of protein-based materials and biomimetic peptides for developing controlled drug delivery systems and advancing biomedical applications.

Figure 1.

Schematic representation of biomimetic peptides discussed in this review. (Figure created using BioRender.com).

2. Design and production of biomimetic peptide conjugates by genetic engineering

Producing biomimetic peptides can be achieved through several synthetic methods, each with unique capabilities and limitations. Examples include solid-phase peptide synthesis (SPPS) (Merrifield, 1963), liquid-phase peptide synthesis (LPPS) (Anderson, 1960), native chemical ligation (NCL) (Dawson et al., 1994), and enzyme-catalyzed peptide synthesis (Stepanov, 1996). These techniques have been thoroughly reviewed in other works (Ullmann & Jakubke, 1999; Nilsson et al., 2005). This review focuses on the production of biomimetic peptides using genetic engineering techniques, which offer advantages in precision, scalability, and functional customization (Lin et al., 2024).

Genetic engineering is distinct because it uses recombinant DNA technology to produce biomimetic peptides directly conjugated to proteins within host cells. Unlike other methods that require post-synthesis chemical conjugation, genetic engineering streamlines production by embedding the desired recombinant peptide-protein gene sequence directly into bacterial, yeast, or mammalian cells, which then express precisely structured conjugates with reduced variability and increased scalability (Lin et al., 2024).

Genetic engineering stands out for several reasons. Firstly, it allows for highly specific control over peptide sequences and their linkage to target proteins. This control enables fine-tuning of the structure and function, such as targeting specific peptide locations on proteins or incorporating flexible linker regions for optimal folding and biological activity (Gupta et al., 2022; Manissorn et al., 2023). Achieving this level of precision with other synthesis methods is more challenging due to the random nature of chemical conjugation or adsorption, which can interfere with bioactivity or release profiles.

Another advantage of genetic engineering is scalability. Once the gene encoding the desired peptide-protein conjugate is integrated into an expression system, it is possible to produce large quantities of uniform, high-purity biomimetic peptide conjugates (Domingo-Espín et al., 2011; Tripathi & Shrivastava, 2019). This scalability is particularly beneficial for therapeutic applications, where consistency is essential. Optimizing expression hosts, such as using bacteria for simpler peptide structures or mammalian cells for complex post-translational modifications, enhances production efficiency and batch consistency, which is essential for clinical translation (Tripathi & Shrivastava, 2019).

Finally, genetic engineering supports the integration of complex functionalities. For instance, with recombinant DNA, researchers can produce multifunctional conjugates that combine multiple biomimetic peptides, each targeting specific binding sites or functionalities (Xu et al., 2022). This enables the creation of conjugates with tailored release kinetics, enhanced binding affinity, and responsiveness to specific biological triggers, thereby expanding the therapeutic and functional potential of biomimetic peptides in controlled release systems (Tu & Tirrell, 2004).

2.1. Recombinant DNA technology in biomimetic peptide conjugates design

Recombinant DNA technology is a powerful method that combines DNA from different sources to engineer cells to produce specific biomolecules. By introducing genetic sequences that encode both biomimetic peptides and therapeutic proteins into host cells, recombinant DNA technology enables the direct intracellular synthesis of these conjugates.

Genes encoding the biomimetic peptide and the therapeutic protein are synthesized and then inserted into a plasmid—a circular DNA molecule that can replicate within a host cell (Figure 2). The plasmid serves as a vector to deliver the genes into host cells. Linkers or specific sequences may be included to ensure that the peptide and protein maintain the correct structure and activity when fused (Chen et al., 2013). The recombinant plasmid is then introduced into a host cell, such as Escherichia coli, yeast, or mammalian cells. These cells then use their own machinery to produce the peptide-protein conjugates. Each host system offers unique advantages: bacterial cells provide rapid, high-yield production, while mammalian cells are capable of performing complex post-translational modifications (Schütz et al., 2023).

Figure 2.

Example plasmid for production of biomimetic peptide conjugates in bacteria. RBS: ribosome-binding site; MBP: maltose-binding protein; TEV: tobacco etch virus protease site; ori: origin of replication; ROP: repressor of primer.

Unlike chemical synthesis methods, which often require additional steps to conjugate the peptide to the protein, recombinant DNA technology enables cells to produce the entire biomimetic peptide-protein conjugates in a single step, making it more suitable for scaling up. This intracellular synthesis reduces the need for additional modifications and minimizes product variability.

2.2. Designing recombinant DNA constructs

To produce biomimetic peptide-protein fusions, plasmid vectors are carefully engineered with specific gene sequences, enabling the design of constructs that fulfill precise therapeutic needs. Plasmid vectors serve as vehicles to deliver synthetic DNA sequences that code for biomimetic peptides and their target proteins into host cells. By introducing custom-designed DNA into plasmids, researchers can direct host cells to synthesize fusion proteins with controlled properties for therapeutic applications.

Based on the example plasmid provided in Figure 2, we will briefly describe the roles and design considerations for each component of the plasmid, as summarized in Table 1.

Table 1.

Example plasmid component for production of biomimetic peptide-protein conjugates.

Plasmid component	Role	Example
Promoter	Expression levels vary depending on the type of promoter used. By selecting strong or inducible promoters, researchers can precisely control the production levels of biomimetic peptide-protein fusions in host cells.	T7 promoter (Studier & Moffatt, 1986) CMV promoter (Thomsen et al., 1984)
Regulatory elements	LacI and LacO play important roles as regulatory elements within the lac operon. LacI encodes the Lac repressor protein, which binds to the LacO operator sequence to block transcription. In the absence of an inducer, such as isopropyl β-D-1-thiogalactopyranoside (IPTG), the Lac repressor prevents RNA polymerase from initiating transcription at the promoter. When IPTG is added, it binds to the Lac repressor, causing a conformational change that releases the Lac repressor from the operator, allowing transcription to proceed. This enables precise control of gene expression by allowing transcription only when IPTG or a similar inducer is present.	LacI LacO (Dyson et al., 2004)
Ribosome-binding site (RBS)	RBS serves as the binding site for the ribosome on the transcribed mRNA, enabling the ribosome to position itself correctly to initiate protein synthesis.	Shine-Dalgarno sequence (Kozak, 1987) Kozak sequence (Kozak, 1989)
Signal peptide sequence (optional)	The signal peptide helps direct the fusion peptide-protein product to a specific cellular location, improving protein folding, stability, and functionality. This process is particularly beneficial for proteins that require secretion, exposure on the cell membrane, or translational modification at specific cellular locations.	PelB leader sequence (Singh et al., 2013)
Purification tag	The purification tag facilitates purification of peptide-protein fusions from bacterial lysates, which is essential for obtaining a high-purity product for downstream applications. The tag can be attached to either the N- or C-terminus of the peptide-protein conjugates. MBP and GST tags are known to enhance the expression and folding of the attached protein.	6xHis tag (Hochuli et al., 1988) GST tag (Smith & Johnson, 1988) MBP tag (di Guan et al., 1988)
Protease recognition site	The protease site allows precise control over the cleavage site of proteins, thereby facilitating the removal of purification tags from peptide-protein conjugates.	Tobacco Etch Virus (TEV) protease cleavage site (Parks et al., 1994) Factor Xa cleavage site (Markmeyer et al., 1990; Jenny et al., 2003) Enterokinase (EK) site (Light & Janska, 1989) Thrombin cleavage site (Jenny et al., 2003)
Sequence encoding biomimetic peptide	The DNA sequence encodes for the biomimetic peptide, and it can be customized to incorporate specific binding motifs or sequences that interact with native proteins in the host system or during its application.	Elastin-like motif (Hu et al., 2015; Wang et al., 2019; Liang et al., 2021) Silk fibroin-mimetic peptide (Yu et al., 2022; Hashimoto et al., 2023; Promsuk et al., 2024) Collagen-mimetic peptide (Chattopadhyay et al., 2016)
Flexible linker	The linker allows the biomimetic peptide to move independently of the therapeutic protein, reducing steric hindrance and ensuring proper folding and function (Chen et al., 2013).	3x[GGGGS] (Trinh et al., 2004) 3x[EAAAK] (Takamatsu et al., 1990)
Sequence encoding bioactive protein	The sequence encodes for therapeutic proteins or other proteins that are useful for bioassays.	Growth factor Fluorescent protein Enzyme Antibody fragment Antimicrobial protein
Termination sequence	The sequence ensures that transcription stops at the end of the gene, providing a clean termination and improving mRNA stability.	T7 terminator (Calvopina-Chavez et al., 2022)
Selection marker	The selection marker allows for the selection of bacteria that have successfully taken up the plasmid, which is essential for ensuring that all cells are capable of peptide production.	Ampicillin-resistance gene (Bolivar et al., 1977) Kanamycin-resistance gene
Origin of replication (Ori)	Allows replication of the plasmid DNA once taken up into host cells.	pBR332 Ori (Bolivar et al., 1977)

2.3. Peptide-protein conjugate expression in host cells

Selecting an appropriate host system for producing peptide-protein conjugates is crucial and largely depends on the protein structural complexity, required yield, and necessary post-translational modifications (PTMs) for functionality. Host cells can be prokaryotic cells or eukaryotic cells. The most common prokaryotic host is E. coli, while eukaryotic hosts include yeasts, insect cells, and mammalian cells (Amann et al., 2019).

Bacterial systems are preferred for expressing simpler peptides and proteins that do not require complex PTMs. Escherichia coli is widely used due to its rapid growth rate, high protein yield, cost-effectiveness, and ease of genetic manipulation. Escherichia coli is particularly suited for peptides that lack disulfide bonds or glycosylation, as bacteria do not possess the cellular machinery to perform eukaryotic PTMs. Moreover, complex protein folding can be problematic, and high expression levels may result in the formation of inclusion bodies (Bhatwa et al., 2021; Manissorn et al., 2023).

Yeast systems offer a combination of the scalability and cost-effectiveness of bacterial systems with the capacity to perform eukaryotic PTMs, such as glycosylation, though with slight structural differences from mammalian systems. Yeasts are suitable for expressing more complex peptides or proteins that require proper folding and PTMs for activity. Pichia pastoris is particularly favored for producing therapeutic proteins at high yields (Shrivastava et al., 2023). Although yeast systems can handle more complex proteins, the glycosylation patterns that they produce differ from those in mammalian cells, potentially affecting bioactivity and therapeutic efficacy (Varki, 2022). Furthermore, peptides with highly complex modifications may not achieve optimal functionality.

Insect cell systems, such as Sf9 and Sf21 cells from Spodoptera frugiperda and High Five cells from Trichoplusia ni, are increasingly favored for expressing complex peptides and proteins, particularly those requiring PTMs that are challenging to achieve in bacterial or yeast systems (Liu et al., 2013). Insect cells can be grown in high-density cultures and in suspension, making them scalable for industrial applications. Baculovirus expression vector systems (BEVS) are commonly employed with insect cells, enabling high levels of target protein production (Liu et al., 2013). However, insect cell cultures are more costly and require more complex media and culture conditions than bacterial or yeast systems, although they are still generally more cost-effective than mammalian cell systems.

Mammalian systems are optimal for expressing highly complex peptides and proteins that require accurate human-like PTMs, such as proper glycosylation patterns, phosphorylation, acetylation, and correct disulfide bond formation. These systems are essential for producing proteins with structures and functions that closely mimic those of the human body, making them ideal for therapeutic applications where bioactivity and safety are critical. However, mammalian cell cultures are more costly, have longer growth cycles, and are generally more challenging to scale than bacterial or yeast systems. The requirement for sterile conditions and specialized media further increases production costs. Nonetheless, these systems are invaluable for peptide-protein conjugates whose function and stability depend on precise PTMs.

2.4. Applications of recombinant DNA technology for biomimetic peptide conjugate production

Recombinant DNA technology facilitates the innovative production of biomimetic peptide conjugates, allowing for the creation of larger, functionally integrated, and scalable constructs that are difficult to achieve through chemical synthesis methods. For example, collagen-mimetic peptides (CMPs) are typically limited to fewer than 50 amino acids when synthesized chemically due to the high cost and challenges associated with large-scale production (Xu & Kirchner, 2021). However, in a study by Jin et al. (2014), CMPs were fused with the minor coat protein of the M13 bacteriophage using recombinant DNA technology (Jin et al., 2014). The viral particles, additionally tagged with fluorescent imaging agents, selectively bound to disrupted collagen structures in lung cancer tissue, enabling precise visualization of pathological sites. This method enabled the integration of a 424-amino-acid viral protein with a CMP sequence of 6–21 amino acids, a scale that is difficult to achieve with chemical synthesis alone, demonstrating the scalability and versatility of the recombinant DNA method in the production of such conjugates.

In another compelling application, Promsuk et al. (2024) used genetic engineering to produce green fluorescent protein (GFP) conjugated with silk fibroin-mimetic peptides (SFMPs) in a series of tandem repeats (Promsuk et al., 2024). The length of the SFMP sequence, ranging from 1 to 7 repeats of the hexapeptide unit (Gly-Ala-Gly-Ala-Gly-Ser), controlled the release profile of GFP from a silk fibroin hydrogel due to specific interactions between the SFMPs and the silk fibroin matrix. This fine-tuning of release rates by SFMP length suggests potential applications in wound healing, where therapeutic proteins, such as fibroblast growth factor (FGF) or epidermal growth factor (EGF), could replace GFP, enabling sustained delivery tailored to promote optimal tissue regeneration.

Another significant example involves elastin-like polypeptides (ELPs), which consist of extensive repeats of the pentapeptide GVGVP, mimicking the elastomeric properties of natural elastin. Using recombinant DNA technology, ELP constructs with up to 251 tandem repeats have been generated, resulting in polypeptides with tunable thermal responsiveness and biocompatibility (McPherson et al., 1996). ELPs have been conjugated to therapeutic proteins for various biomedical applications, including anti-cancer therapeutics [e.g. IFN (Hu et al., 2015; Wang et al., 2019; Liang et al., 2021), anti-FLT3 (Park et al., 2020), anti-CD99 (Vaikari et al., 2020), DRA (Manzari et al., 2019)], diabetes management [e.g. FGF-21 (Gilroy et al., 2018)], and wound healing [e.g. vRAGE (Kang et al., 2021)]. By enabling precise control over ELP length and conjugation with diverse therapeutic agents, recombinant DNA technology allows researchers to create multifunctional bioconjugates designed for controlled release, targeted delivery, and high stability in vivo.

These examples highlight the versatility of recombinant DNA technology, enabling the customization of biomimetic peptide conjugates in ways that surpass chemical synthesis in both complexity and functional design. This approach allows researchers to design highly specific, scalable, and bioactive peptide-based therapeutics suitable for a range of advanced medical applications.

3. Collagen-mimetic peptide conjugates for drug delivery

Collagens are structural protein components of the extracellular matrix (ECM) in various tissues, such as skin, tendons, bones, cartilage, blood vessels, and ligaments. In addition to providing structural support, these proteins also provide signals to neighboring cells, which are crucial for angiogenesis, hemostasis, tissue repair, and biomineralization (San Antonio et al., 2020). In biomaterials, collagens have increasingly gained attention from scientists because they are biocompatible, biodegradable, non-immunogenic, and can be reconstituted into fibrous structures in various forms of biomedical materials, such as hydrogels, scaffolds, films, and particles. To date, more than 20 different types of collagens have been identified. Collagens exist as triple helices, which are rigid, rope-like protein conformations. The triple-helical regions are composed of proline-rich tripeptides, specifically, Gly-X-Y repeats. The X and Y positions are typically occupied by proline and hydroxyproline, respectively. Hydroxyproline is necessary for the formation of intramolecular hydrogen bonds, which stabilize the triple-helical conformation (Figure 3(A)) (Gelse et al., 2003). Since glycine has the smallest side chain among amino acids, glycine side chains are packed inside the helix, while the side chains of proline and hydroxyproline face outward (Abou Neel et al., 2013).

Figure 3.

Collagen structure and its interactions with collagen-mimetic peptides (CMPs). (A) The hierarchical structure of collagen. (B) A collagen scaffold containing a bioactive protein tagged with a CMP. The CMP interacts with the scaffold’s collagen. (Figure created using BioRender.com).

To form a collagen fibril, the N- and C-termini of a triple-helical procollagen are cleaved by procollagen peptidases. This enzymatic cleavage results in tropocollagen, a building block for collagen fibrils. Tropocollagens self-assemble in a staggered fashion with regular spacing, producing a distinct banding pattern in electron micrographs (Zhu et al., 2018). Tropocollagens are reinforced by covalent crosslinks, including aldol and aldimine bonds, between the staggered molecules. The self-assembled, crosslinked tropocollagens form a long fibrous structure known as a collagen fibril. Multiple collagen fibrils then align to form larger collagen fibers. The hierarchical organization of collagen is the basis for its high tensile strength and flexibility (Bhattacharjee & Bansal, 2005).

The role of collagen in biomaterials extends beyond structural applications to the creation of biocompatible, biodegradable, and non-immunogenic carriers that mimic the natural ECM environment. This versatility has led to its use in various biomedical forms, including hydrogels, scaffolds, films, and particles, for applications in tissue engineering and drug delivery systems.

Early works on collagen as a biomaterial focused on fabricating extracted collagen into various forms, such as bone-regeneration scaffolds, hydrogels, bandages, and tendon repair patches (Abou Neel et al., 2013). Later, collagen was shown to be promising for drug delivery applications (Khan & Khan, 2013). Several studies were conducted using collagen as a drug carrier in drug delivery systems (Figure 3(B)). For example, collagen types I and III were used as carriers to sustain the release of gentamicin (Ruszczak & Friess, 2003; Maczynska et al., 2019). Collagen mini-pellets were developed for interferon release, which, in turn, prolonged serum tumor necrotic factor (TNF) concentrations (Fujioka et al., 1995). These collagen matrices provide a scaffold for drug loading but lack the specificity needed to directly target pathological sites.

Recent advancements in collagen-mimetic peptides (CMPs) address these limitations, as CMPs are designed to hybridize with damaged or denatured collagen found at pathological sites, thereby providing a targeted approach to drug delivery (Table 2). Researchers have developed CMPs that can potentially form triple-helical structures with collagen. The design of CMPs is based on the Gly-X-Y repeat, with some modifications. CMPs are capable of specifically targeting pathological collagen in vivo through triple-helix hybridization (Li et al., 2012). The specificity of CMPs to interact with pathological collagen enables imaging and delivery of therapeutics at the pathological sites. Additionally, CMP conjugates can be used as bioactive building blocks to construct smart biomaterials through self-assembly (Luo & Kiick, 2017). For example, fluorescently labeled CMPs can target solid tumors through triple-helix hybridization, as the high MMP-9 activity in tumor tissue leads to the exposure of denatured collagen (Li et al., 2012, 2013). The triple-helix hybridization has also been exploited to confer collagen-targeting properties to nanoparticles. Collagen-targeted nanoparticles can sustain the release of nucleic acids from collagen platforms (Urello et al., 2014, 2016). Furthermore, CMPs were synthesized as (Gly-X-Y)₇-tagged growth factors. The CMPs can bind strongly to natural collagen in wound tissue, offering the potential for long-term, wound-specific attachment and controlled release of growth factors. Therefore, CMP-tagged effector molecules could be useful for wound healing in burn and diabetic patients (Chattopadhyay et al., 2012). Chadttophadhyay et al. explored a method to promote wound healing by conjugating the cytoactive factor, substance P, to a CMP that selectively binds to damaged collagen, serving as a vehicle to localize therapeutic agents precisely where they are needed in the wound. In wound healing tests on diabetic mice (a model for impaired wound healing), wounds treated with CMP-conjugated substance P healed more effectively, demonstrating enhanced re-epithelialization and reduced inflammation compared to controls (Chattopadhyay et al., 2016). Another study by the same research group utilized the conjugation of CMP to a TGF-β-clustering peptide (Chattopadhyay et al., 2022). The conjugates anchor to exposed collagen in wounds and enhance local activation by binding and clustering TGF-β, a key factor in tissue repair. This approach resulted in improved cellular responses, including enhanced migration and proliferation of cells essential for wound healing. In diabetic mouse models with impaired healing, treatment with this peptide significantly improved healing rates, supporting better re-epithelialization and reducing scar formation compared to controls. This bioconjugate strategy demonstrated the potential for developing targeted and sustained-release systems for cytoactive factors to improve chronic wound care.

Table 2.

Collagen-mimetic peptides.

Biomimetic peptide	Biomimetic sequence	Bioactive conjugate	Application	References
Streptococcal collagen-like2 (Scl2) protein conjugated to glycosaminoglycan (GAG)-binding protein	(Gly-Xaa-Yaa)n	GAG-binding protein	hMSCs to chondrocyte differentiation (hydrogel)	Parmar et al., 2015
Collagen mimetic peptide (CMP) conjugated to hyaluronic acid (HA)	GPO)8-CG-RGDS	HA	BMSC to cartilage differentiation (hydrogel)	Ren et al., 2020
6-Pattern collagen-like peptide conjugated to silane	Ac-Lys-[Pro-Hyp-Gly]6-Lys	Silane	MSC encapsulation for cartilage tissue engineering (hydrogel)	Valot et al., 2021
CMP conjugated to substance P	(ProProGly)7	Substance P	Wound healing (peptide-protein conjugate)	Chattopadhyay et al., 2016
CMP conjugated to TGF-β receptor ligand	(ProProGly)7	TGF-β receptor ligand (LTGKNFPMFHRN)	Wound healing (peptide-peptide conjugate)	Chattopadhyay et al., 2022

Furthermore, Parmar et al. designed MMP7-degradable hydrogels using the collagen-mimetic Scl2 protein, which contains the repeating unit (Gly-Xaa-Yaa)_n conjugated with glycosaminoglycan (GAG)-binding peptides (Parmar et al., 2015). The presence of GAG-binding peptides significantly increased the chondrogenic differentiation of Scl2 hydrogels impregnated with human mesenchymal stem cells (hMSCs). This biodegradable hydrogel could be used for articular cartilage regeneration and other regenerative medicine applications. Ren et al. synthesized a hyaluronic acid (HA)-based hydrogel conjugated with CMPs, (GPO)₈-CG-RGDS, and crosslinked it with MMP-sensitive peptides (Ren et al., 2020). This biodegradable HA hydrogel induced chondrogenic differentiation of bone marrow stem cells (BMSCs). Moreover, Valot et al. developed a biomimetic chemical hydrogel based on silylated CMPs, which was used for cartilage tissue engineering (Valot et al., 2021).

However, integrating CMPs into controlled-release systems presents several challenges. CMPs are relatively short peptides, often limited to fewer than 50 amino acids when synthesized chemically. This size restriction not only increases production costs but also limits scalability. Recombinant DNA technology has addressed this limitation, enabling the efficient and affordable production of larger, more complex CMP conjugates. For instance, Jin et al. (2014) developed CMPs conjugated to the minor coat proteins of the M13 bacteriophage, creating a hybrid structure capable of selectively binding to disrupted collagen in lung cancer tissue (Jin et al., 2014). This approach exemplifies how recombinant CMPs overcome size limitations, enabling the creation of multifunctional constructs with applications in targeted imaging and therapy.

4. Elastin-like polypeptide conjugates in therapeutic protein delivery

Elastin is a resilient connective-tissue protein found in the ECM that provides tissues with elasticity and strength. This protein is abundant in blood vessels, lungs, elastic ligaments, tendons, and skin (Wang et al., 2021). It plays essential role in wound healing, scar repair, and the progression of cardiovascular diseases (Baumann et al., 2021; Wang et al., 2021). The precursor of elastin is tropoelastin, which is encoded by 36 exons, some of which are subject to alternative splicing (Kanta, 2016). Tropoelastin contains two major domains: hydrophobic and hydrophilic domains. Elastin consists of 850–870 amino acids (72–74 kDa) and contains several hydrophobic domains alternating with hydrophilic domains. The hydrophobic domains are rich in valine (V), glycine (G), and proline (P) (Bailey, 1978). These hydrophobic domains consist of repeating units of 3–6 peptides, such as GVGVA, GVGVP, GGVP, and GVAP. The repetitive hydrophobic domains tend to form β-turns, which are essential for elasticity (Kanta, 2016). The hydrophilic crosslink domains are rich in alanine (A) and lysine (K) that tend to form α-helices. The α-helices contribute to the insolubility of elastin (Figure 4(A)) (Boyd et al., 1991; Kanta, 2016). Lysine residues in the hydrophilic domains can be crosslinked by lysyl oxidase, forming a continuous, long, and flexible elastin fiber.

Figure 4.

Elastin structure and elastin-like peptide-based protein carriers. (A) Elastin contains alternating hydrophobic and hydrophilic domains that tend to adopt β-turns and α-helices, respectively. Lysine residues can be crosslinked by lysyl oxidase to form elastin fibers. (B) The ELP structure mimics elastin repeats and undergoes reversible self-assembly to form micelles for the controlled release of a protein. (Figure created using BioRender.com).

Elastin-like polypeptides (ELPs), a soluble form of elastin mimetics, contain pentapeptide repeats—valine-proline-glycine-X-glycine (VPGXG)—in which X can be any natural amino acid except proline (Urry, 1984; Roberts et al., 2015). The repeats are derived from the hydrophobic domain of tropoelastin. The number of repeats and variations in residue X can be used to manipulate the ELP properties, such as the transition temperature, electromagnetic responses, ionic strength, and pH sensitivity. ELPs are thermally responsive and can self-assemble as the temperature increases (Figure 4(B)). ELPs are typically designed as diblocks. One block is hydrophobic, while the other is hydrophilic. Below the critical micelle temperature (CMT), ELP diblocks remain soluble. In contrast, above the CMT, the hydrophobic domain selectively desolvates, converting the diblock ELPs into amphiphiles and driving their self-assembly into micelles (Saha et al., 2020). Chemical, enzymatic, and physical crosslinking methods have been used to prepare ELP-based hydrogels. For example, lysine-containing ELPs were rapidly crosslinked to form hydrogels using tris(hydroxymethyl)phosphine propionic acid. Fibroblasts embedded in ELP hydrogels survived and remained viable after 3 days in vitro, suggesting potential applications in tissue engineering (Lim et al., 2007). Furthermore, ELP nanoparticles were produced from diblock copolymers with distinct X residues and block lengths. This approach enabled control over the protein transition temperature during micelle formation (Dreher et al., 2008).

ELP micelles have been used in various applications, including the delivery of a wide array of bioactive agents, ranging from chemotherapeutics, such as doxorubicin (Bidwell et al., 2007) and geldanamycin (Chen et al., 2011) to growth factors, such as BMP2 and BMP14 (Bessa et al., 2010). Recently, researchers have used ELP carriers to enhance the therapeutic efficacy of clinically relevant drugs by tuning their ­polymeric structure. Callahan et al. designed pH-responsive polypeptide micelles containing ELP repeating units, VPG[VG₇A₈]G and VPG[VH₄]G. The multiple-stimuli-sensitive ELP-based particles were designed to dissociate in the presence of a low-pH environment within solid tumors (Callahan et al., 2012). Their work provides valuable insights for ­designing strategies to enhance tumor targeting and ­imaging agents.

Furthermore, ELP hydrogels can be formed from the same basic structure of VPGXG repeating units. Additional crosslinking methods have been successfully examined using cysteine residues incorporated into the polymer backbone. These residues form disulfide bonds between strands, reducing the weight percentage of polymer required to form hydrogels (Asai et al., 2012). Wang et al. developed a hybrid ELP-PEG hydrogel to improve the optical properties of ELP hydrogels. The ELP-PEG hydrogel enhanced optical transparency and promoted fibroblast viability. This method has potential for use in cell imaging or the study of cell-matrix interactions (Wang et al., 2014).

Drug release systems were developed based on ELPs, including the use of thermally triggered phase transitions for the delivery of glucagon-like peptides in diabetes treatment (Table 3). These peptides were expressed as fusions with ELPs and released in response to a temperature change. The released glucagon-like peptides were shown to reduce blood glucose levels for up to 5 days after administration (Amiram et al., 2013). Similarly, injectable gel-forming ELP-curcumin conjugates for delivering anti-inflammatory molecules to the sciatic nerve showed promise in limiting the systemic dissemination of the anti-inflammatory drug, with satisfactory release sustained over 4 days (Sinclair et al., 2013).

Table 3.

Elastin-like polypeptides.

Biomimetic peptide	Biomimetic sequence	Bioactive conjugate	Application	References
Elastin-like polypeptide (ELP) with Cys substitution	(VPGXG)n where X = Cys, Ala, or Val n = 6, 10, or 16	FITC-BSA	Sustained release of BSA as model (hydrogel)	Asai et al., 2012
Cell-binding-ELP	TVYAVTGRGDSPASSAA [(VPGIG)2VPGKG(VPGIG)2]3VP	RGD sequence	Study of cell–matrix interactions (hydrogel)	Wang et al., 2014
ELP conjugated to glucagon-like peptide 1 (GLP-1)	(GAGVPGGGVP)60GY and (GVGVP)120GWP	Glucagon-like peptide 1 (GLP-1)	Sustained release of GLP-1 to control blood glucose Peptide-peptide conjugate/self-assembled aggregates	Amiram et al., 2013
ELP conjugated to curcumin	MSKGPG[VPGXG]L = 60, 80, 160 WPC with X = V/I/E [1:3:1]	Curcumin carbamate	Anti-inflammation	Sinclair et al., 2013
ELP conjugated to IFN-α	(VPGXG)90 X = A/G/V [2:3:5]	IFN-α	Enhanced pharmacokinetics, tumor accumulation, and antitumor efficacy	Hu et al., 2015
ELP conjugated to anti-FTL3	G(VPGAG)192Y	Anti-FTL3	Extended pharmacokinetic half-life for myeloid leukemia treatment (nanoparticles)	Park et al., 2020
ELP conjugated to anti-CD99	G(VPGAG)192Y	Anti-CD99	Extended pharmacokinetic half-life for myeloid leukemia treatment (nanoworms)	Vaikari et al., 2020
ELP conjugated to DR5 agonist (DRA)	(VPGVG)120	DRA	Tumor growth inhibition (insoluble coacervate)	Manzari et al., 2019
ELP conjugated to FGF-21	(VPGXG)120 X = V/A/ [4:1]	FGF-21	Sustained anti-diabetic action (insoluble coacervate)	Gilroy et al., 2018
ELP conjugated to vRGAE	V40C2, where V = VPGVG and C = VPGVGVPGVGVPGCGVPGVGVPGVG	vRGAE	Diabatic wound treatment (insoluble coacervate)	Kang et al., 2021

ELPs have also been conjugated to therapeutic proteins for various biomedical applications, including anti-cancer therapeutics [e.g. IFN (Hu et al., 2015; Wang et al., 2019; Liang et al., 2021), anti-FLT3 (Park et al., 2020), anti-CD99 (Vaikari et al., 2020), and DRA (Manzari et al., 2019)]. This demonstrates the utility of biomimetic polypeptide conjugation and responsiveness to local conditions, such as temperature and enzymes, in enhancing drug delivery efficacy, particularly in solid tumors, where targeted and sustained delivery can improve therapeutic outcomes. In addition to cancer therapy, ELPs have been conjugated to FGF-21 for diabetes management (Gilroy et al., 2018). Gilroy et al. examined the fusion of FGF-21 with a thermally responsive ELP to develop an injectable depot for diabetes treatment (Gilroy et al., 2018). The ELP-FGF-21 complex forms a stable depot under physiological conditions, enabling sustained FGF-21 release over extended periods and resulting in prolonged anti-diabetic effects in vivo. This approach minimizes the need for frequent injections and demonstrates potential as a long-lasting therapeutic for managing diabetes by ­maintaining FGF-21 levels within therapeutic concentrations. Furthermore, Kang et al. developed self-assembling ELP-vRAGE fusion protein coacervates as inhibitors of advanced glycation end-products (AGEs), which are associated with impaired wound healing in diabetes (Kang et al., 2021). The ELP-vRAGE coacervates competitively inhibit AGE formation, reduce inflammation, and promote wound healing in diabetic models. This method offers a novel therapeutic approach by addressing AGE-related complications in diabetic wounds and demonstrates that ELP coacervates enhance the wound healing environment by acting as targeted inhibitors of harmful biochemical pathways.

Despite their numerous benefits, ELPs exhibit certain limitations. One concern is their relative mechanical instability compared to natural elastin, particularly in vivo, where maintaining structural integrity over extended periods is crucial for effective drug delivery. The use of crosslinking agents—such as lysine-containing ELPs crosslinked by tris(hydroxymethyl)phosphine—can enhance stability but may introduce unwanted side effects or immunogenicity in therapeutic applications (Adams et al., 2009). Furthermore, although ELPs show promise in the encapsulation and delivery of drugs, such as doxorubicin and therapeutic proteins, such as BMPs, their tendency to aggregate at physiological temperatures may limit their suitability for temperature-sensitive biological drugs, necessitating extensive formulation optimization.

Moreover, the tenability of ELPs in response to pH and ionic strength broadens their versatility but also poses challenges for achieving consistent release profiles, particularly in heterogeneous environments, such as tumor tissue, where pH may vary significantly (Wike-Hooley et al., 1985). The pH-responsive ELP micelles designed by Callahan et al. exemplify efforts to address this challenge by incorporating histidine residues (VPG[VH4]G) to enable selective release in acidic tumor environments (Callahan et al., 2012). While promising, these complex designs may hinder large-scale production and require rigorous in vivo testing to confirm their efficacy and safety.

5. Silkworm silk fibroin-mimetic peptide conjugates for drug delivery

Silk fibroin (SF) is a naturally derived biological polymer extracted from silkworm cocoons. Silk fibroin from silkworms (Bombyx mori) is the most extensively studied form of SF for biomedical applications. Silk cocoons are spun from fibers that consist of two major proteins: silk fibroin and silk sericin. Sericin is a glue-like protein with a molecular weight ranging from 20 to 400 kDa (Kunz et al., 2016). The complex hydrophilic structure consists of 18 amino acids with polar groups, including amine and amide, that are available for crosslinking, polymerization, and integration with other polymers (Kunz et al., 2016). Silk fibroin is a fibrous protein composed of a heavy chain (391 kDa) and a light chain (26 kDa) linked by a single disulfide bond at the C-terminus and a p25 glycoprotein in a molar ratio of 6:6:1 (Holland et al., 2019). The primary structure of the heavy chain of silk fibroin consists of 12 hydrophobic repetitive domains flanked by 11 hydrophilic non-repetitive linker domains. The amino acid sequence of the repetitive domains is often divided into four types of motifs: GAGAGS, GAGAGY, GAGAGV, and GAGAGVGY (Fu et al., 2009; Holland et al., 2019). The hydrophobic repetitive domains fold and bond together via hydrogen bonds, van der Waals forces, and hydrophobic interactions to form anti-parallel β-sheet structures (Figure 5). The β-sheets contribute to most of the crystalline regions, whereas the non-repetitive linkers are amorphous in SF materials (Thongnuek et al., 2023).

Figure 5.

Silk fibroin structure and its interactions with silk fibroin-mimetic peptides (SFMPs) within a scaffold. (A) The structure of silk fibroin, featuring crystalline anti-parallel β-sheets and amorphous linkers. The sequence of SF repeating units typically used for biomimetic design is shown. (B) The interactions of SFMP-tagged bioactive molecules with the crystalline domains of SF are key to sustaining the release of these molecules. (Figure created using BioRender.com).

A significant advantage of SF-based materials in drug delivery is their versatility across various forms, including films, disks, porous scaffolds, hydrogels, and nanoparticles, enabling varied release kinetics and degradation rates (Duangpakdee et al., 2021; Manissorn et al., 2021; Watchararot et al., 2021; Sapudom et al., 2023). This adaptability facilitates the controlled release of small-molecule drugs, proteins, and nucleic acids. For example, SF-based hydrogels have been shown to stabilize antibodies and enable their gradual release without compromising protein functionality (Guziewicz et al., 2013; Zhang et al., 2017). However, natural silk protein alone may be limited in certain applications due to its relatively slow degradation rate, which may not align with the pharmacokinetics of some drugs. Additionally, traditional drug loading methods, such as immersing hydrogels or scaffolds in drug solutions, can result in burst release.

To address these limitations, SF-mimetic peptides, typically based on the GAGAGS repeat, have been developed to mimic SF β-sheet-forming domains and to be used as fusion tags, structural enhancers, or components in synthetic polymers. For example, Yu et al. developed silk fibroin repeats (GAGAGS)_n as a fusion tag (SF-tag) to enhance the expression of nanobodies in Escherichia coli (Yu et al., 2022). Their work showcases the protein-stabilizing properties of SF in bacterial systems, which is promising for scaling up peptide production. Huang et al. demonstrated a strategy to generate rubber-like protein hydrogels using photo-crosslinkable resilin-like blocks (GGRPSDSYGAPGGGN)_n and silk-like blocks (GAGAGS)_n (Huang et al., 2021). Dinerman et al. studied the swelling behavior of a physically crosslinked hydrogel, which was composed of silk-elastin-like protein (SELP) containing silk blocks (GAGAGS)_n and elastin blocks (GVGVP)_n (Dinerman et al., 2002). SELPs combined semi-crystalline silk blocks and elastomeric elastin blocks. The engineered SELP hydrogels could be used for the delivery of genes and viruses (Huang et al., 2015; Wani et al., 2022). To facilitate scale-up production, SELP was produced using recombinant DNA technology (Fernández-Colino et al., 2014). Gustafson et al. demonstrated the potential of SELP-815K hydrogels as an injectable controlled gene delivery platform for the treatment of head and neck cancer via gene-directed enzyme prodrug therapy (Gustafson et al., 2010). However, while SELP-based hydrogels can be thermally responsive and injectable, maintaining consistent behavior across various physiological environments remains challenging, as thermosensitivity may vary depending on tissue type or pathological conditions. Huang et al. synthesized thermoresponsive copolymers of recombinant resilin-silk (RS), which were useful for fabricating hydrogel materials (Huang et al., 2017). Hashimoto et al. conjugated the GRGDS peptide with the silk fibroin repeat (GAGAGS) and immobilized it on an SF film. The conjugated GRGDS(GAGAGS) enhanced the adhesion and elongation of NIH(3T3) cells on the surface of the SF film (Table 4) (Hashimoto et al., 2023).

Table 4.

Silk fibroin mimetic peptides.

Biomimetic peptide	Biomimetic sequence	Bioactive conjugate	Application	References
SFMP conjugated to RGD	GAGAGS	RGD sequence (GRGDS)	Cell adhesion improvement (film)	Hashimoto et al., 2023
SFMP conjugated to GFP	(GAGAGS)n=0–6	GFP	Tunable release rate of GFP (hydrogel)	Promsuk et al., 2024
SFMP conjugated to nanobody	(GAGAGS)n=1–5	Anti-HE4 nanobodies	Improving nanobody yield.	Yu et al., 2022
SFMP conjugated to ELP	GAGAGS in silk blocks GXGVP in Elastin blocks with various ratios of silk to elastin blocks	Doxorubicin	Cancer cell inhibition (nanoparticles)	Xia et al., 2014
SFMP conjugated to resilin	[(GGRPSDSYGAPGGGN)4(GAGAGS)4]5	–	Nanoparticles/fibers/coacervates anticipated for drug delivery	Huang et al., 2017

The development of SFMPs with varied GAGAGS repeat lengths represents a novel approach to controlling release kinetics and improving mechanical properties. Promsuk et al. demonstrated that longer GAGAGS repeats in SFMPs prolonged protein release from SF hydrogels and improved mechanical stability, highlighting their potential for diverse therapeutic applications (Promsuk et al., 2024). However, further studies are needed to validate these results in vivo, as changes in peptide length and structure may affect immunogenicity and degradation profiles, potentially leading to inconsistent therapeutic outcomes.

Although GAGAGS-based peptides show great promise in drug delivery, several challenges in production, design, and application merit deeper exploration. First, the SFMP unique amino acid sequence, GAGAGS, which mimics the repetitive motifs of native silk fibroin, plays a crucial role in recreating β-sheet crystallinity. However, producing SFMPs in E. coli poses significant metabolic challenges due to their high glycine and alanine content. Escherichia coli cells have limited pools of these amino acids, which means that prolonged or scaled-up expression of SFMPs could deplete cellular resources, potentially slowing growth and protein production. This challenge may be overcome by using metabolically engineered E. coli that have an elevated glycyl-tRNA pool, as has been done in the production of recombinant spider silk (Xia et al., 2010).

Additionally, the high hydrophobicity of the GAGAGS motif complicates protein folding and solubility, increasing the risk of aggregation or inclusion body formation. Moreover, the hydrophobic nature of GAGAGS restricts available synthesis options. Chemical synthesis techniques struggle with repetitive hydrophobic sequences, as these segments are prone to aggregation, which makes achieving high-purity yields difficult.

From a materials perspective, hydrophobicity and β-sheet crystallinity are essential for SFMP stability, yet they limit the versatility of SFMP-based drug delivery. While β-sheets provide mechanical strength and slow-release properties, they may also hinder SFMP biodegradation and inhibit rapid therapeutic release needed for short-term treatments. Therefore, SFMP engineering should prioritize achieving a balance by integrating hydrophilic motifs or linker domains to enhance both solubility and degradation control. Studies, such as that by Promsuk et al., which showed that tailoring SFMP length can adjust release rates and material strength, offer insights into creating SFMPs better suited to clinical needs (Promsuk et al., 2024).

6. Spider silk-based peptide conjugates for drug delivery

Spider silk is one of the most well-studied biomaterials due to its outstanding mechanical properties. Spiders can produce more than one type of silk (Eisoldt et al., 2011). Among the different types of spider silks, draglines from the golden orb weaver (Nephila clavipes) and the garden spider (Araneus diadematus) are the most well-characterized (Scheibel, 2004). The two major proteins of draglines from N. clavipes are major ampullate spidroins (MaSp1 and MaSp2), which have highly repetitive core sequences. Another well-characterized spider silk is the dragline silk from the garden spider (A. diadematus), which is composed mainly of two large proteins, namely A. diadematus fibroin 3 (ADF3) and A. diadematus fibroin 4 (ADF4) (Scheibel, 2004). Similar to SF, MaSp and ADF consist of crystalline β-sheet structures interspersed with amorphous linkers. The crystalline β-sheets are formed by interactions among the polyalanine motifs, A_n or (GA)_n. The amorphous linkers, which serve as elastic domains, are formed by the motifs (GGX)_n or GPGXX (where X represents tyrosine, leucine, or glutamine) (Figure 6(A)) (Eisoldt et al., 2011). The molecular weights of dragline silk proteins are ∼250–350 kDa (Hagn, 2012). Recently, an engineered spider silk protein called eADF4 was developed. The genetically engineered protein consists of 16 repeats of a module named C (GSSAAAAAAAASGPGGYG PENQGPSGPGGYGPGGP), and the protein is sometimes called eADF4 (C16). The eADF4 (C16) mimics the consensus sequence of the dragline silk protein ADF4 from the garden spider (Figure 6(A)). Due to its biocompatibility, biodegradability, and non-toxicity, the eADF4 (C16) mimetic peptide has been used in various biomedical applications. Moreover, it can be fabricated into various forms, such as hydrogels, capsules, particles, and films, for drug delivery (Blüm et al., 2014; Dou et al., 2019; Weiss et al., 2020; Laomeephol et al., 2021; Yazawa et al., 2022).

Figure 6.

Schematic representation of recombinant spider silk and its functional applications. (A) The structure of eADF4 (C16) and spidroins. (B) Interactions of eADF4 (C16)-tagged bioactive molecules within a spider silk-based hydrogel, highlighting its potential for biomedical applications. (Figure created with BioRender.com).

Besides its use in structural materials, recombinant spider silk protein has recently been explored for drug and protein delivery (Figure 6(B) and Table 5). Schacht et al. investigated the hydrogel formation of the spider silk protein eADF4 (C16), functionalized with fluorescein, using different protein concentrations and chemical crosslinking (Schacht & Scheibel, 2011). An increase in the eADF4 (C16) concentration reduced the gelation time. Chemical crosslinking stabilized the hydrogel and increased its mechanical strength. The eADF4 (C16) functionalized with fluorescein exhibited low shear and elastic moduli in the resulting hydrogel. Variations in the physical properties of the hydrogel could influence the pharmacokinetics of encapsulated molecules. However, crosslinking agents may also affect biocompatibility, potentially increasing the risk of immune reactions or cytotoxicity. It has been shown that different types of crosslinking agents affect immune cells in distinct ways (Sapudom et al., 2023). Kumari et al. utilized recombinant spider silk eADF4 (C16) hydrogel for the sustained release of biological molecules, including bovine serum albumin (BSA), horseradish peroxidase (HRP), and lysozyme (LYS) (Kumari et al., 2018). The researchers employed various loading methods, including direct loading, diffusion loading, and particle-in-gel loading, and subsequently observed the release profiles. The release of the three biologics was most effectively prolonged from hours to weeks when loaded using the particle-in-gel method. Arndt et al. employed the N-terminal domain of recombinant spider silk proteins, which formed amyloid-like fibrils during gelation and preserved the functionality of fusion proteins (Arndt et al., 2022). Luo et al. synthesized a reversible hydrogel using copolymers of the C-terminus of spider silk and resilin, which exhibited thermoresponsive and pH-responsive properties (Luo et al., 2018). This multistimuli-responsive hydrogel may prove useful for protein or drug delivery; however, it may face challenges in maintaining consistent response thresholds within complex physiological environments. Rigorous in vivo testing is necessary to ensure that these materials perform reliably across diverse conditions.

Table 5.

Spider silk-mimetic peptides.

Biomimetic peptide	Biomimetic sequence	Bioactive conjugate	Application	References
eADF4 (C16)	(GSSAAAAAAASGPGPGGYP PENQGPSGPGGYGPGGP)16	–	Hydrogel/nanofibrils for drug delivery	Schacht & Scheibel, 2011
eADF4 (C16) conjugated to DNA aptamer	(GSSAAAAAAASGPGPGGYP PENQGPSGPGGYGPGGP)16	DNA aptamer	Capture and release thrombin from silk-based surfaces	Humenik et al., 2020
Spidroin N-terminal domain (NT)	GSGNSHTTPWTNPGLAENFMNSFMQ GLSSMPGFTASQLDDMSTIAQSMVQS IQSLAAQGRTSPNKLQALNMAFASSMA EIAASEEGGGSLSTKTSSIASAMSNAFLQ TTGVVNQPFINEITQLVSMFAQAGMNDVS	GFP Purine nucleoside phosphorylase	Protein immobilization and controlled release (hydrogel)	Arndt et al., 2022
MaSP1-based silk peptide conjugated to polylysine	SGRGGLGGQGAGAAAAAGGAGQGGYGGLGSQGT	15xLysine	pDNA complexation to deliver to cancer cells (nanoparticles)	Numata et al., 2012
MaSP1-based silk peptide conjugated to hepticindin or human neutrophil defensins	(SGRGGLGGQGAGAAAAAGGAGQGGYGGLGSQGT)6	Hepticidin or human neutrophil defensins	Antibiotic applications (film)	Gomes et al., 2011, 2012

In terms of spider silk conjugates, Humenik et al. utilized functionalized DNA aptamer-spider silk eADF4 (C16) nanohydrogels in an immobilization technique for the bio-selective and reversible binding of proteins (Humenik et al., 2020). In this study, anti-thrombin aptamers were chemically conjugated to eADF4 (C16). The aptamer-eADF4 (C16) conjugates selectively bound human thrombin within the nanohydrogel. The release of active thrombin was triggered by conformational changes in anti-thrombin aptamers induced by complementary DNA sequences. The modified DNA aptamer-conjugated eADF4 (C16) spider silk demonstrated the potential of biomimetic peptides for immobilizing bioactive molecules within a hydrogel. This approach shows promise for applications requiring high specificity, such as anticoagulant delivery. However, conjugating aptamers to spider silk proteins complicates synthesis, increases costs, and heightens sensitivity to environmental conditions. Overcoming these barriers will require the development of cost-effective synthesis and stabilization techniques, as the thermal and chemical sensitivity of aptamers can limit the practicality of such systems in real-world conditions.

Numata et al. engineered silk-based nanocomplexes using genetic engineering to conjugate spider silk units with polylysine domains, which can form complexes with plasmid DNA and tumor-homing peptides (Numata et al., 2012). The nanocomplexes were designed to deliver DNA specifically to cancer cells, utilizing silk fibroin as a biocompatible carrier with enhanced stability and tumor-targeting capabilities. Tumor-homing peptides enhanced the selectivity of the nanocomplexes, leading to increased uptake by tumor cells compared to healthy cells. The results demonstrated effective gene expression within tumor tissues, highlighting the potential for precise gene therapy in cancer treatment.

Gomes and colleagues focus on creating antimicrobial spider silk by genetically fuzing it with antibiotic proteins, including Hepcidin and human neutrophil defensins (Gomes et al., 2011, 2012). They demonstrated that spider silk peptides conjugated with antibiotic proteins represent a promising biomaterial for preventing infection and supporting tissue repair. The engineered antimicrobial spider silk proteins in these studies demonstrate strong potential for drug delivery, particularly in applications requiring local antimicrobial action, such as wound healing and tissue engineering. However, translating this approach into broader drug delivery applications will require addressing several challenges, including controlling the release rate of drugs embedded within the silk, ensuring the stability and activity of the fused antimicrobial agents over time, and optimizing production scalability for clinical use.

7. Fibrin-based peptides

Fibrin is a natural biopolymer that plays a critical role in blood coagulation, wound healing, angiogenesis, and inflammation (Laurens et al., 2006). It is naturally derived from the cleavage of fibrinogen by thrombin during blood coagulation (Undas & Ariëns, 2011). Fibrin acts as a natural scaffold for wound closure and the regeneration of wounded tissue.

Fibrinogen is a hexameric homodimer glycoprotein (340 kDa) composed of two sets of three polypeptide chains (Aα-, Bβ-, and γ-), each with a fibrinopeptide cleavage site (Figure 7(A)) (Pieters & Wolberg, 2019). The dimers are linked by 29 disulfide bonds. Fibrin formation occurs through thrombin cleavage at the fibrinopeptide cleavage sites on the N-termini of the Aα and Bβ chains, resulting in the formation of unstable protofibrils. Covalent crosslinking of fibrins by Factor XIII results in a stable structure and enhances elastic properties (Figure 7(B)).

Figure 7.

From fibrinogen to fibrin to film: Structural and biochemical transitions. (A) The structure of fibrinogen, composed of two sets of three polypeptide chains (Aα, Bβ, and γ) with fibrinopeptide cleavage sites located at N-termini. These cleavage sites are essential for the activation of fibrin formation. (B) Fibrin formation mediated by factor XIII crosslinking, which stabilizes the fibrin network and facilitates its conversion into a stable film. (Figure created with BioRender.com).

Fibrin-based biomaterials offer substantial promise for drug delivery and tissue engineering due to their inherent biocompatibility, hemostatic properties, and ability to form a stable, crosslinked matrix that mimics the natural wound-healing process in the body (Spicer & Mikos, 2010; Noori et al., 2017; Jarrell et al., 2021). Additionally, fibrin can be fabricated into various forms, including hydrogels, scaffolds, and films (Pereira et al., 2023; Zhou et al., 2023; Zou et al., 2023). Several studies have utilized fibrin as a copolymer and incorporated fibrin-binding peptides for tissue engineering applications. Gila-Vilchez et al. developed a composite hydrogel by reacting fibrin with short peptides (Fmoc-FF and Fmoc-RGD), which enhanced the mechanical properties of the hydrogel (Gila-Vilchez et al., 2023). The composite hydrogel exhibited increased storage and shear moduli compared to a hydrogel made from pure fibrin. Furthermore, the RGD moiety improved cell viability and ex vivo biocompatibility of the hydrogel. Brinkmann et al. fabricated a fibrin-based hydrogel composed of fibrinogen and a fibrin-binding peptide (sequence: PCDYYGTCLD or Tn7). The Tn7 peptides significantly enhanced Young’s modulus and mechanical stiffness of the hydrogel. The peptides also showed no toxicity, making them suitable for tissue regeneration and cell-based therapies (Brinkmann et al., 2023). However, questions remain regarding the balance between stability and biodegradability. Excessive stabilization can hinder the timely degradation required for tissue remodeling, potentially resulting in prolonged inflammatory responses. Additionally, fibrin-mimetic peptides have been used for various purposes. For example, Sanborn et al. developed a 20-residue peptide from the γ-chain of fibrin (CTIGEGQQHHLGGAKQAGDV), conjugated with four-armed branched PEG, to create an in situ hydrogel (Table 6). The components of the peptide-PEG hydrogel consisted of calcium-loaded liposomes, thrombin, and factor XIII, which were thermally triggered to generate elastic gel formation at 37 °C. The created hydrogel is useful for drug and gene delivery (Sanborn et al., 2002). Nandi et al. synthesized platelet-like hydrogel microparticles (microgels), which were functionalized with fibrin-mimetic peptide knob B (AHRPYAAK). The microgel increased clot density in vitro and decreased bleeding in a rodent trauma model in vivo (Nandi et al., 2021).

Table 6.

Fibrin-mimetic peptides.

Biomimetic peptide	Biomimetic sequence	Bioactive conjugate	Application	References
Factor XIII crosslinking site of fibrin with Cys addition	CTIGEGQQHHLGGAKQAGDV	–	Blood-coagulation-mimicking hydrogel proposed for drug and gene delivery	Sanborn et al., 2002
Fibrin knob-B mimetics conjugated to pNIPAm microgel	AHRPYAAK	–	Increase clot density and decrease bleeding.	Nandi et al., 2021

While fibrin-based materials are generally well-tolerated, immunogenic responses remain a concern, especially in cases involving allogenic or xenogenic sources. These immune responses could compromise drug delivery efficacy and patient safety. Furthermore, the natural role of fibrin in hemostasis can pose a double-edged sword: while beneficial for wound closure, the potential for unintended clot formation in non-hemostatic contexts is a significant risk factor that requires careful consideration in therapeutic designs. It is also worth noting that no research has attempted to conjugate fibrin-based peptides to a therapeutic protein for targeted or controlled delivery.

8. Keratin-based peptides

Keratin is a natural polymer obtained from various biological sources, such as hair, skin, wool, nails, and feathers (Yang et al., 2017; Abascal & Regan, 2018; Rajabi et al., 2020), many of which can be agriculturally enriched, while others are wastes from esthetic grooming. Keratin is an insoluble fibrous protein that contains a high level of cysteine (Abascal & Regan, 2018). Keratins can be classified into two types: α-keratin (40–68 kDa) and β-keratin (10–22 kDa) (Wang et al., 2016). α-keratin, found in mammals, is a helical protein with coiled-coil structures. The primary structure of keratins contains an amino acid repeating unit of (X₁X₂X₃X₄X₅X₆X₇)_n, where X represents any amino acid (Figure 8(A)) (Wang et al., 2016; Yang et al., 2017). The heptapeptide coiled-coil structures of keratins, such as EVSALEK, KVSALKE, EIAALEK, KIAALKE, VAALEKE, and VAALKEK, have been utilized in the creation of protein-based biomaterials (Yang et al., 2017). In contrast, β-keratin forms β-sheets, which are the major structural components of bird and reptile keratins (Figure 8(B)). Additionally, keratins can be classified into soft and hard keratins, which differ in their sulfur and lipid content (Wang et al., 2016). Furthermore, the coiled-coil structures of α-keratin and many fibrous proteins are valuable for protein and drug delivery applications, including coiled-coil-based hydrogels, coiled-coil polymer hybrids, and coiled-coil nanostructures (Figure 8(A)) (Utterström et al., 2021).

Figure 8.

Schematic representation of keratin structures and their biomaterial applications. (A) The coiled-coil structure of α-keratin and its fabrication into various biomaterial forms, highlighting its mechanical properties and versatility. (B) The pleated β-sheet structure of β-keratin, emphasizing its role in providing rigidity and structural stability. (Figure created with BioRender.com).

Keratins have been used in many aspects of tissue engineering and the protein can be fabricated into various biomaterial formats, such as scaffolds, films, hydrogels, and fibers (Rajabi et al., 2020). Ham et al. fabricated a keratin hydrogel using a combination of keratose (KOS) and kerateine (KTN) to control the release rate of recombinant human insulin-like growth factor 1 (rhIGF-1). Increasing KOS led to an increase in the release rate of rhIGF-1 from the KOS/KTN hydrogel (Ham et al., 2016). They also showed that a high amount of KTN and a low amount of KOS increased the elastic modulus and mechanical strength. Interestingly, KTN:KOS ratios appeared not to affect porosity or cell viability (Ham et al., 2016). Kan et al. studied the solubilization process of recombinant keratin from human hair keratin 37 and keratin 81 and elucidated the protein structure with the goal of improving wound-healing ability (Kan et al., 2020). Similarly, Gao et al. investigated the dermal wound healing capability of recombinant keratin derived from human hair keratin 37 and keratin 81. The authors found that keratin nanoparticles enhanced cell proliferation and migration in vitro and improved wound healing in vivo (Gao et al., 2019).

The tunability of drug release remains limited in comparison to other synthetic and biological polymers. For long-term release, keratin degradation may not be sufficiently predictable, as it is influenced by factors, such as the presence of enzymes and oxidative conditions in vivo. Moreover, keratin biodegradation can be either too rapid or too slow for drug delivery purposes, particularly in physiological environments where enzymatic activities are variable. As a result, modifying keratin to achieve an ideal degradation profile without compromising its bioactivity and compatibility remains a central challenge in keratin-based drug delivery systems.

Additionally, keratin extraction methods vary widely, and the quality of extracted keratin is often inconsistent due to differences in source material and purification processes. The harsh chemical treatments (e.g. reduction and oxidation) required to solubilize keratin can disrupt its natural structure and bioactivity, potentially affecting the reproducibility and mechanical properties of keratin-based biomaterials. In particular, optimizing cysteine content for disulfide bond formation is crucial, as this dictates the material stability, elasticity, and degradation rate. While a high cysteine content can enhance mechanical stability, it also increases stiffness, which may limit the flexibility required for certain drug delivery applications. Despite these challenges, keratin-based materials offer a sustainable and potentially customizable platform for drug delivery, with ongoing advancements in recombinant technology and hybridization with synthetic polymers improving control over their properties. Similar to fibrin, keratin-mimetic peptides are relatively under-researched, especially compared to other biomimetic peptides, such as collagen-, elastin-, and silk-based peptides. This limited research interest could be attributed to several factors. The high cysteine content in keratin, which forms dense disulfide bonds, makes it more rigid and challenging to process. This complexity poses a major hurdle for the design and synthesis of keratin-mimetic peptides that accurately replicate the properties of natural keratin.

9. Resilin-based peptides

Resilin is an elastic protein mostly found in insect cuticles, such as those of locusts, dragonflies, cicadas, and cockroaches (Su et al., 2014; Michels et al., 2016). The resilin protein sequence was first identified in Drosophila melanogaster as the CG15290 gene product. The sequence consists of a signal peptide at the N-terminus and three different exons (exons 1–3). Exons 1 and 3 encode hydrophilic repeating units, which include 18 repeats of GGRPSDSYGAPGGGN and 11 repeats of GYSGGRPGGQDLG, respectively. Exon 2 encodes a chitin-binding domain (Figure 9) (Su et al., 2014). Exons 1 and 3 are rich in the amino acids proline (P) and glycine (G), which provide structural flexibility (Tamburro et al., 2010). Resilin has β-turn and β-spiral conformations, which are similar to those of other elastomeric proteins (Tamburro et al., 2010; Qin et al., 2012; Yang et al., 2017). Furthermore, resilin shows rubber-like characteristics, exhibiting two outstanding material properties: high resilience and high elasticity (Weis-Fogh, 1960; Gosline et al., 2002).

Figure 9.

Schematic representation of the structure of resilin from Drosophila melanogaster. (Figure created using BioRender.com).

Resilin-like polypeptides (RLPs) and resilin-mimetic peptides have been studied in various reports (Table 7). In 2020, Balu et al. described the physicochemical properties and gel-cell interactions of a resilin hydrogel by modifying the structure of a resilin-like polypeptide (RLP) using graphene oxide (GO). The incorporation of GO into the Rec1/GO hybrid hydrogel resulted in increased hydrophilicity, resilience, swelling ratio, micropore size, amorphous domain size, and cell proliferation. The incorporation of GO reduced the crosslinking density, mass fractal cluster size, compressive elastic modulus, and cell-inert characteristics (Balu et al., 2020). Balancing reduced mechanical strength with enhanced biocompatibility remains challenging, as high resilience may be advantageous for flexibility but can compromise structural stability and load-bearing capabilities compared to synthetic polymers or other natural biopolymers. This tradeoff is especially important in tissue engineering applications, where structural stability is critical.

Table 7.

Resilin mimetic peptides.

Biomimetic peptide	Biomimetic sequence	Bioactive conjugate	Application	References
Resilin-like polypeptide (RLP)	(GGRPSDSYGAPGGGN)18	–	Incorporated with graphene oxide for gel–cell interaction study	Truong et al., 2011; Balu et al., 2020
RLP with DOPA substitution	[(GGRPSDSYGAPGGGNY)4]n=4 or 8 Y can be modified to DOPA by enzyme	–	Hydrogel potentially for drug carrier	Zhu et al., 2023
Resilin-elastin fusion polypeptides	MGKKKPVSDTYGA (PGGGNGGRPSDTYGA)4-(Elastin blocks)	GRGDSP peptide	Cell adhesive nanofibers	Bracalello et al., 2019

Truong et al. constructed a resilin-mimetic peptide, Rec1-resilin, from the first exon of the D. melanogaster CG15920 gene and investigated the effects of water vapor sorption and hydration on the molecular chain dynamics and viscoelastic properties of resilin hydrogel (Truong et al., 2011). While these findings underscore the adaptability of resilin, achieving optimal hydration in vivo is complex, as dynamic fluctuations in tissue fluid levels can unpredictably impact mechanical performance. Further research is needed to improve control over resilin hydration and moisture retention, which could enable more consistent material properties and expand its applications in drug delivery systems. Homma et al. developed a simple method for the expression of recombinant resilin-like protein using Brevibacillus choshinensis (Homma et al., 2023). Zhu et al. established a method to incorporate the noncanonical amino acid DOPA into resilin-like protein (RLP). The DOPA content and molecular length of RLP played important roles in determining the viscoelastic and self-healing properties of the resilin hydrogel (Zhu et al., 2023). Moreover, Bracalello and collaborators designed chimeric resilin-elastin polypeptides to enhance cell-binding ability. The fabricated nanofibers from these chimeric resilin-elastin polypeptides were non-cytotoxic and could be used in biomedical devices (Bracalello et al., 2019).

The use of resilin-based peptides in drug delivery remains a relatively underexplored field. Resilin research has primarily focused on utilizing its unique elastic and resilient properties for tissue engineering and the development of biomaterials that mimic soft tissues. Studies on resilin-like polypeptides (RLPs) and resilin-mimetic peptides have often highlighted their high resilience, elasticity, and biocompatibility, but applications have largely focused on fields, such as mechanical tissue support, adaptive hydrogels, and dynamic biointerfaces.

10. Conclusion, challenges, and future perspectives

This review highlights significant advancements in using biomimetic peptides to control the release of bioactive molecules from various biopolymer platforms. These biopolymers—including collagen, elastin, silk fibroin, spider silk, fibrin, keratin, and resilin—offer unique properties, such as biocompatibility, biodegradability, and mechanical strength, making them ideal for various biomedical applications. Biomimetic peptides, inspired by the natural sequences and structures of these biopolymers, have been pivotal in improving the functionality and specificity of drug delivery systems and tissue engineering scaffolds.

Collagen-mimetic peptides (CMPs) enhance targeted delivery and facilitate smart biomaterial design by replicating the natural interactions within the ECM, thereby aiding in tissue repair and angiogenesis. Innovations in CMPs enable precise control over bioactive molecule release, improving therapeutic outcomes. Elastin-like polypeptides (ELPs), derived from the elastin amino acid sequence, demonstrate thermally responsive self-assembly, which facilitates controlled release mechanisms. The ability of ELPs to form micelles and hydrogels in response to temperature changes makes them suitable for delivering a variety of therapeutic agents. Silk fibroin (SF), characterized by repetitive β-sheet structures, provides stability and strength. Biomimetic peptides based on SF primary structures, such as the GAGAGS motif, have been used to enhance drug delivery and tissue regeneration by providing controlled and sustained release of bioactive molecules. Spider silk proteins, particularly recombinant versions, such as eADF4(C16), mimic the natural silk mechanical properties and flexibility. These engineered proteins can form various structures, including hydrogels and nanohydrogels, enabling the sustained and controlled release of therapeutic agents. Fibrin, a crucial component in wound healing and blood coagulation, forms robust scaffolds for tissue regeneration. Biomimetic peptides that mimic fibrin-binding sites enhance the stability and mechanical properties of fibrin-based hydrogels, making them more suitable for controlled drug delivery. Keratin, owing to its structural integrity, is employed in hydrogels and scaffolds to promote wound healing and cell proliferation. Biomimetic approaches utilizing keratin natural sequences have shown the potential in enhancing the delivery of growth factors and other bioactive molecules. Resilin, recognized for its exceptional elasticity and resilience, has inspired the design of biomimetic peptides capable of forming elastic and durable hydrogels. Resilin-like peptides have been utilized to create materials that replicate the mechanical properties of natural resilin, improving the controlled release and mechanical stability of systems for bioactive molecule delivery.

While utilizing biomimetic peptides to enhance biopolymer platforms for the controlled release of bioactive molecules holds significant promise, several challenges must be addressed to fully realize their potential in clinical and biomedical applications. These challenges range from design and synthesis to functionalization and regulatory approval. The design of biomimetic peptides that precisely replicate the natural functions and interactions of native proteins is inherently complex. Achieving precise sequence, proper folding, and functionality can be particularly challenging for peptides required to perform specific biological activities. The screening of appropriate refolding conditions can be particularly daunting (Manissorn et al., 2023). In terms of stability, biomimetic peptides are often susceptible to degradation by proteases within the biological environment. Maintaining the stability of these peptides without compromising their functionality is crucial. This process often necessitates chemical modifications or the incorporation of non-natural amino acids, which can significantly complicate synthesis. Recent advancements in peptoid synthesis may offer promising solutions to address the stability challenges associated with biomimetic peptides (Zuckermann, 2011). The synthesis of high-fidelity biomimetic peptides is often both expensive and labor-intensive. The development of cost-effective and scalable production methods, such as advanced recombinant DNA technologies or peptide synthesis techniques, is imperative for practical applications.

Furthermore, the integration of biomimetic peptides into biopolymer platforms, while preserving their bioactivity and functionality, presents significant challenges. The conjugation process must be meticulously controlled to prevent denaturation or loss of bioactivity. Attaining precise control over the release kinetics of bioactive molecules from biopolymer platforms remains a significant challenge. The interactions between biomimetic peptides and the biopolymer matrix must be precisely tuned to achieve the desired release profile, which can vary depending on the specific application and the therapeutic molecules involved (Promsuk et al., 2024).

Future research may emphasize the development of multifunctional biomimetic peptides. For instance, conjugating peptides with bio-orthogonal groups may facilitate real-time tracking, while incorporating bioactive motifs could improve cell-targeting capabilities. These modifications could enhance both functionality and monitoring capabilities during delivery, providing researchers and clinicians with greater control over therapeutic outcomes. Given the versatility and unique properties of less-explored peptides, such as keratin- and resilin-mimetic peptides, further investigation into these materials could address existing gaps in tissue engineering and drug delivery applications. The development of these peptides could benefit from advanced tools in protein engineering and biopolymer synthesis to replicate native functions and resolve stability challenges in vivo. A significant gap exists in the study of the immunogenicity of biomimetic peptides, particularly those containing modified amino acids or chimeric peptides, such as resilin-elastin fusions. The minimization of immune responses remains an ongoing challenge in the clinical translation of biomimetic peptides. Studies investigating the immunogenicity of biomimetic peptides and strategies to mitigate adverse responses—such as PEGylation, glycosylation, or host-specific sequence design—will be essential, particularly for long-term therapies. As these materials progress into in vivo applications, ensuring minimal immune responses will facilitate their transition to widespread clinical use.

The future of biomimetic peptide-enhanced biopolymer platforms for the controlled release of bioactive molecules appears highly promising. Ongoing research into biomimetic peptides is expected to focus on enhancing their specificity for targeting particular tissues or cellular microenvironments. Such advancements are anticipated to enhance the efficacy of drug delivery systems, reduce side effects, and maximize therapeutic benefits. The ability to tailor biomimetic peptide-enhanced biopolymer platforms to individual patient needs, such as varying release kinetics for diabetes patients, has the potential to revolutionize personalized medicine. Customized scaffolds and delivery systems are expected to optimize treatments for specific conditions, thereby improving patient outcomes. In the context of bioprocess engineering, developing scalable and cost-effective methods for producing biomimetic peptides and their biopolymer conjugates will be essential for their widespread clinical adoption. Advances in recombinant technology, peptide synthesis, and bioreactor systems are likely to play a pivotal role in achieving this goal.

Funding Statement

PT is supported by The Asahi Glass Foundation (RES_66_111_2100_011). He is also supported by the NSRF via the Program Management Unit for Human Resources & Institutional Development, Research, and Innovation (B38G670008). PT is funded by National Research Council of Thailand (NRCT) and Chulalongkorn University (N42A670600). JM is supported by Ratchadapisek Somphot Fund for Postdoctoral Fellowship, Chulalongkorn University. JP was supported by the Science Achievement Scholarship of Thailand (SAST), the 90th Anniversary of Chulalongkorn University Scholarship, and the Research Assistantship Funding from Faculty of Science, Chulalongkorn University (RAF_2562_012). KW was supported by the Chulalongkorn University grant to the Center of Excellence for Molecular Biology and Genomics of Shrimp, Chulalongkorn University grant to the Center of Excellence in Molecular Crop. This research is funded by Thailand Science Research and Innovation Fund Chulalongkorn University (BCG_FF_68_294_2100_040).

Author contributions

Juthatip Manissorn and Jaturong Promsuk drafted the original manuscript. Juthatip Manissorn and Jaturong Promsuk also took part in the conception, analysis, and interpretation of the data in the manuscript. Kittikhun Wangkanont and Peerapat Thongnuek conceptualized, designed, and critically reviewed the manuscript. Also, Kittikhun Wangkanont and Peerapat Thongnuek acquired fundings and supervised the manuscript drafting. All authors agree to be accountable for all aspects of the work.

Disclosure statement

No potential conflict of interest was reported by the author(s).

Data availability statement

Data sharing is not applicable to this article as no new data were created or analyzed in this study.